U.S. patent number 8,226,815 [Application Number 12/569,159] was granted by the patent office on 2012-07-24 for small volume in vitro sensor and methods of making.
This patent grant is currently assigned to Abbott Diabetes Care Inc.. Invention is credited to Fredric C. Colman, Benjamin J. Feldman, Jeffery V. Funderburk, Adam Heller, Ephraim Heller, Rajesh Krishnan, Fei Mao, Joseph A. Vivolo.
United States Patent |
8,226,815 |
Feldman , et al. |
July 24, 2012 |
Small volume in vitro sensor and methods of making
Abstract
A sensor utilizing a non-leachable or diffusible redox mediator
is described. The sensor includes a sample chamber to hold a sample
in electrolytic contact with a working electrode, and in at least
some instances, the sensor also contains a non-leachable or a
diffusible second electron transfer agent. The sensor and/or the
methods used produce a sensor signal in response to the analyte
that can be distinguished from a background signal caused by the
mediator. The invention can be used to determine the concentration
of a biomolecule, such as glucose or lactate, in a biological
fluid, such as blood or serum, using techniques such as coulometry,
amperometry, and potentiometry. An enzyme capable of catalyzing the
electrooxidation or electroreduction of the biomolecule is
typically provided as a second electron transfer agent.
Inventors: |
Feldman; Benjamin J. (Oakland,
CA), Heller; Adam (Austin, TX), Heller; Ephraim
(Piedmont, CA), Mao; Fei (Fremont, CA), Vivolo; Joseph
A. (San Francisco, CA), Funderburk; Jeffery V.
(Stevenson Ranch, CA), Colman; Fredric C. (Woodside, CA),
Krishnan; Rajesh (San Leandro, CA) |
Assignee: |
Abbott Diabetes Care Inc.
(Alameda, CA)
|
Family
ID: |
27379577 |
Appl.
No.: |
12/569,159 |
Filed: |
September 29, 2009 |
Prior Publication Data
|
|
|
|
Document
Identifier |
Publication Date |
|
US 20100018877 A1 |
Jan 28, 2010 |
|
Related U.S. Patent Documents
|
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
Issue Date |
|
|
11734979 |
Apr 13, 2007 |
|
|
|
|
10662081 |
Sep 12, 2003 |
7225535 |
|
|
|
09594285 |
Jun 15, 2000 |
6618934 |
|
|
|
09295962 |
Apr 21, 1999 |
6338790 |
|
|
|
60103627 |
Oct 8, 1998 |
|
|
|
|
60105773 |
Oct 8, 1998 |
|
|
|
|
Current U.S.
Class: |
205/792; 205/775;
205/777.5 |
Current CPC
Class: |
G01N
27/3271 (20130101); G01N 27/3274 (20130101); C12Q
1/001 (20130101); C12Q 1/006 (20130101); G01N
27/3272 (20130101); C12Q 1/32 (20130101); Y10T
29/49004 (20150115); Y10T 29/4921 (20150115); Y10T
29/49147 (20150115); Y10T 29/49208 (20150115); Y10T
29/49007 (20150115); Y10T 29/4913 (20150115); Y10T
29/49128 (20150115); Y10T 29/49002 (20150115); Y10T
29/49126 (20150115); Y10T 29/49794 (20150115); Y10T
156/1052 (20150115); Y10T 29/49155 (20150115); Y10T
29/49204 (20150115) |
Current International
Class: |
G01N
27/327 (20060101) |
Field of
Search: |
;204/403.01-403.15
;205/777.5,778,792,775 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
|
|
|
|
|
|
|
227029 |
|
Sep 1985 |
|
DD |
|
2903216 |
|
Aug 1979 |
|
DE |
|
3934299 |
|
Oct 1990 |
|
DE |
|
4234553 |
|
Apr 1993 |
|
DE |
|
29720299 |
|
Jan 1998 |
|
DE |
|
0010375 |
|
Apr 1980 |
|
EP |
|
0026995 |
|
Apr 1981 |
|
EP |
|
0048090 |
|
Mar 1982 |
|
EP |
|
0078636 |
|
May 1983 |
|
EP |
|
0080304 |
|
Jun 1983 |
|
EP |
|
0096288 |
|
Dec 1983 |
|
EP |
|
0136362 |
|
Apr 1984 |
|
EP |
|
0125139 |
|
Nov 1984 |
|
EP |
|
0127958 |
|
Dec 1984 |
|
EP |
|
0170375 |
|
Feb 1986 |
|
EP |
|
0177743 |
|
Apr 1986 |
|
EP |
|
0184909 |
|
Jun 1986 |
|
EP |
|
0206218 |
|
Dec 1986 |
|
EP |
|
0230472 |
|
Aug 1987 |
|
EP |
|
0241309 |
|
Oct 1987 |
|
EP |
|
0245073 |
|
Nov 1987 |
|
EP |
|
0255291 |
|
Feb 1988 |
|
EP |
|
0278647 |
|
Aug 1988 |
|
EP |
|
0286084 |
|
Oct 1988 |
|
EP |
|
0359831 |
|
Mar 1990 |
|
EP |
|
0368209 |
|
May 1990 |
|
EP |
|
0390390 |
|
Oct 1990 |
|
EP |
|
0400918 |
|
Dec 1990 |
|
EP |
|
0453283 |
|
Oct 1991 |
|
EP |
|
0470290 |
|
Feb 1992 |
|
EP |
|
0470649 |
|
Feb 1992 |
|
EP |
|
0537761 |
|
Apr 1993 |
|
EP |
|
0781406 |
|
Jul 1997 |
|
EP |
|
1060707 |
|
Dec 2000 |
|
EP |
|
1318815 |
|
May 1973 |
|
GB |
|
1394171 |
|
May 1975 |
|
GB |
|
1599241 |
|
Sep 1981 |
|
GB |
|
2073891 |
|
Oct 1981 |
|
GB |
|
2154003 |
|
Aug 1985 |
|
GB |
|
2204408 |
|
Nov 1988 |
|
GB |
|
2254436 |
|
Oct 1992 |
|
GB |
|
54-41191 |
|
Apr 1979 |
|
JP |
|
55-10581 |
|
Jan 1980 |
|
JP |
|
55-10583 |
|
Jan 1980 |
|
JP |
|
55-10584 |
|
Jan 1980 |
|
JP |
|
55-12406 |
|
Jan 1980 |
|
JP |
|
56-163447 |
|
Dec 1981 |
|
JP |
|
57-70448 |
|
Apr 1982 |
|
JP |
|
57-98853 |
|
Jun 1982 |
|
JP |
|
58-211646 |
|
Dec 1983 |
|
JP |
|
59-34882 |
|
Feb 1984 |
|
JP |
|
59-67452 |
|
Apr 1984 |
|
JP |
|
59-147249 |
|
Aug 1984 |
|
JP |
|
60-173457 |
|
Sep 1985 |
|
JP |
|
60-173458 |
|
Sep 1985 |
|
JP |
|
60-173459 |
|
Sep 1985 |
|
JP |
|
60-211350 |
|
Oct 1985 |
|
JP |
|
61-2060 |
|
Jan 1986 |
|
JP |
|
62-114747 |
|
May 1987 |
|
JP |
|
62-139629 |
|
Jun 1987 |
|
JP |
|
63-58149 |
|
Mar 1988 |
|
JP |
|
63-128252 |
|
May 1988 |
|
JP |
|
63-139246 |
|
Jun 1988 |
|
JP |
|
63-294799 |
|
Dec 1988 |
|
JP |
|
63-317757 |
|
Dec 1988 |
|
JP |
|
63-317758 |
|
Dec 1988 |
|
JP |
|
64-54345 |
|
Mar 1989 |
|
JP |
|
1-114746 |
|
May 1989 |
|
JP |
|
1-114747 |
|
May 1989 |
|
JP |
|
1-124060 |
|
May 1989 |
|
JP |
|
1-134244 |
|
May 1989 |
|
JP |
|
1-134245 |
|
May 1989 |
|
JP |
|
1-134246 |
|
May 1989 |
|
JP |
|
1-156658 |
|
Jun 1989 |
|
JP |
|
1-291153 |
|
Nov 1989 |
|
JP |
|
02-19758 |
|
Jan 1990 |
|
JP |
|
2-62958 |
|
Mar 1990 |
|
JP |
|
2-120655 |
|
May 1990 |
|
JP |
|
2-245650 |
|
Oct 1990 |
|
JP |
|
2-287145 |
|
Nov 1990 |
|
JP |
|
2-310457 |
|
Dec 1990 |
|
JP |
|
3-26956 |
|
Feb 1991 |
|
JP |
|
3-28752 |
|
Feb 1991 |
|
JP |
|
3-165249 |
|
Jul 1991 |
|
JP |
|
3-202764 |
|
Sep 1991 |
|
JP |
|
03-293556 |
|
Dec 1991 |
|
JP |
|
4-194660 |
|
Jul 1992 |
|
JP |
|
04-264246 |
|
Sep 1992 |
|
JP |
|
5-72171 |
|
Mar 1993 |
|
JP |
|
05-149910 |
|
Jun 1993 |
|
JP |
|
5-196595 |
|
Aug 1993 |
|
JP |
|
06-109688 |
|
Apr 1994 |
|
JP |
|
06-130032 |
|
May 1994 |
|
JP |
|
61-90050 |
|
Jul 1994 |
|
JP |
|
62-85855 |
|
Oct 1994 |
|
JP |
|
7-27734 |
|
Jan 1995 |
|
JP |
|
7-55757 |
|
Mar 1995 |
|
JP |
|
7-72585 |
|
Mar 1995 |
|
JP |
|
7-270373 |
|
Oct 1995 |
|
JP |
|
8-285814 |
|
Nov 1996 |
|
JP |
|
8-285815 |
|
Nov 1996 |
|
JP |
|
8-320304 |
|
Dec 1996 |
|
JP |
|
9-21778 |
|
Jan 1997 |
|
JP |
|
9-101280 |
|
Apr 1997 |
|
JP |
|
09-159642 |
|
Jun 1997 |
|
JP |
|
09-166571 |
|
Jun 1997 |
|
JP |
|
9-264870 |
|
Oct 1997 |
|
JP |
|
9-285459 |
|
Nov 1997 |
|
JP |
|
10-2874 |
|
Jan 1998 |
|
JP |
|
10-170471 |
|
Jun 1998 |
|
JP |
|
1281988 |
|
Jan 1987 |
|
SU |
|
WO 85/05119 |
|
Nov 1985 |
|
WO |
|
WO 86/00513 |
|
Jan 1986 |
|
WO |
|
WO 89/08713 |
|
Sep 1989 |
|
WO |
|
WO 90/05300 |
|
May 1990 |
|
WO |
|
WO 90/05910 |
|
May 1990 |
|
WO |
|
WO 91/01680 |
|
Feb 1991 |
|
WO |
|
WO 91/04704 |
|
Apr 1991 |
|
WO |
|
WO 91/09139 |
|
Jun 1991 |
|
WO |
|
WO 91/15993 |
|
Oct 1991 |
|
WO |
|
WO 92/13271 |
|
Aug 1992 |
|
WO |
|
WO 94/20602 |
|
Sep 1994 |
|
WO |
|
WO 94/27140 |
|
Nov 1994 |
|
WO |
|
WO 95/02817 |
|
Jan 1995 |
|
WO |
|
WO 95/13534 |
|
May 1995 |
|
WO |
|
WO 95/22597 |
|
Aug 1995 |
|
WO |
|
WO 95/28634 |
|
Oct 1995 |
|
WO |
|
WO 96/00614 |
|
Jan 1996 |
|
WO |
|
WO 96/06947 |
|
Mar 1996 |
|
WO |
|
WO 96/30431 |
|
Oct 1996 |
|
WO |
|
WO 96/32635 |
|
Oct 1996 |
|
WO |
|
WO 97/00441 |
|
Jan 1997 |
|
WO |
|
WO 97/02847 |
|
Jan 1997 |
|
WO |
|
WO 97/13870 |
|
Apr 1997 |
|
WO |
|
WO 97/18464 |
|
May 1997 |
|
WO |
|
WO 97/18465 |
|
May 1997 |
|
WO |
|
WO 97/19344 |
|
May 1997 |
|
WO |
|
WO 97/42882 |
|
Nov 1997 |
|
WO |
|
WO 97/42883 |
|
Nov 1997 |
|
WO |
|
WO 97/42886 |
|
Nov 1997 |
|
WO |
|
WO 97/42888 |
|
Nov 1997 |
|
WO |
|
WO 97/43962 |
|
Nov 1997 |
|
WO |
|
WO 98/01208 |
|
Jan 1998 |
|
WO |
|
WO 98/35225 |
|
Aug 1998 |
|
WO |
|
WO 98/35225 |
|
Aug 1998 |
|
WO |
|
WO 98/43073 |
|
Oct 1998 |
|
WO |
|
WO 98/58250 |
|
Dec 1998 |
|
WO |
|
WO 99/08106 |
|
Feb 1999 |
|
WO |
|
WO 99/30152 |
|
Jun 1999 |
|
WO |
|
WO 00/78210 |
|
Dec 2000 |
|
WO |
|
WO 01/64105 |
|
Sep 2001 |
|
WO |
|
WO 01/72220 |
|
Oct 2001 |
|
WO |
|
WO 01/73395 |
|
Oct 2001 |
|
WO |
|
Other References
Batchelor et al., Amperometric Assay for the Ketone body
3-Hydroxybutyrate, Analytica Chimica Acta, 221 (1989) 289-294.
cited by examiner .
JPO English language translation of Tadahisa JP 09-101280
downloaded Apr. 28, 2006. cited by examiner .
Bindra, D.S. et al., "Design and in Vitro Studies of a Needle-Type
Glucose Sensor for Subcutaneous Monitoring", Anal. Chem.,
63(17):1692-1696 (Sep. 1, 1991). cited by other .
Bobbioni-Harsch, E. et al., "Lifespan of subcutaneous glucose
sensors and their performances during dynamic glycaemia changes in
rats," J. Biomed. Eng. 15:457-463 (1993). cited by other .
Bowyer et al., "Electrochemical Measurements in Submicroliter
Volumes", Analytical Chemistry, 64, pp. 459-462 (1992). cited by
other .
Brandt, J. et al., "Covalent attachment of proteins to
polysaccharide carriers by means of benzoquinone," Biochim.
Biophys. Acta, 386(1) (1 page Abstract only) (1975). cited by other
.
Bratten et al. "Micromachining Sensors for Electrochemical
Measurement in Subnanoliter Volumes" Analytical Chemistry, vol. 69,
No. 2, (Jan. 15, 1997). cited by other .
Brownlee, M. et al., "A Glucose-Controlled Insulin-Delivery System:
Semisynthetic Insulin Bound to Lectin", Science,
206(4423):1190-1191 (Dec. 7, 1979). cited by other .
Cass, A.E.G. et al., "Ferrocene-Mediated Enzyme Electrode for
Amperometric Determination of Glucose", Anal. Chem., 56(4):667-671
(Apr. 1984). cited by other .
Cass, A.E.G. et al., "Ferricinum Ion as an Electron Acceptor for
Oxido-Reductases," J. Electroanal. Chem., 190:117-127 (1985). cited
by other .
Cassidy et al., "Novel Electrochemical Device for the Detection of
Cholesterol or Glucose" Analyst, Apr. 1993 vol. 118 pp. 415-418.
cited by other .
Castner, J. F. et al., "Mass Transport and Reaction Kinetic
Parameters Determined Electrochemically for Immobilized Glucose
Oxidase," Biochemisty, 23(10):2203-2210 (1984). cited by other
.
Chen, C.Y. et al., "A Biocompatible Needle-Type Glucose Sensor
Based on Platinum-Electroplated Carbon Electrode", Applied
Biochemistry and Biotechnology, 36:211-226 (1992). cited by other
.
Claremont, D.J. et al., "Biosensors for Continuous In Vivo Glucose
Monitoring", IEEE Engineering in Medicine and Biology Society 10th
Annual International Conference, New Orleans, Louisiana, 3 pp.
(Nov. 4-7, 1988). cited by other .
Clark, L.C. et al., "Differential Anodic Enzyme Polarography for
the Measurement of Glucose", Oxygen Transport to Tissue:
Instrumentation, Methods, and Physiology, 127-133 (1973). cited by
other .
Clark, L.C. et al., "Long-term Stability of Electroenzymatic
Glucose Sensors Implanted in Mice," Trans. Am. Soc. Artif. Intern.
Organs, XXXIV:259-265 (1988). cited by other .
Clarke, W. L., et al., "Evaluating Clinical Accuracy of Systems for
Self-Monitoring of Blood Glucose," Diabetes Care, 10(5):622-628
(Sep.-Oct. 1987). cited by other .
Creager et al. ("Self-assembled monolayers and enzyme electrodes:
progress, problems and prospects", Analytica Chimica Acta
307:277-289 (1995)). cited by other .
Csoregi, E. et al., "Design, Characterization, and One-Point in
Vivo Calibration of a Subcutaneously Implanted Glucose Electrode,"
Anal. Chem. 66(19):3131-3138 (Oct. 1, 1994). cited by other .
Csoregi, E. et al., "On-Line Glucose Monitoring by Using
Microdialysis Sampling and Amperometric Detection Based on "Wired"
Glucose Oxidase in Carbon Paste," Mikrochim. Acta. 121:31-40
(1995). cited by other .
Csoregi, E. et al., "Design and Optimization of a Selective
Subcutaneously Implantable Glucose Electrode Based on "Wired"
Glucose Oxidase," Anal. Chem. 67(7):1240-1244 (Apr. 1, 1995). cited
by other .
Darahazi and Tokuda, "Cyclic voltammetry for reversible
redox-electrode reactions I thin-layer cells with closely separated
working an auxiliary electrodes of the same size", J. Electroanaly.
Chem, 264, p. 77-89, (1989). cited by other .
Davis, G., "Electrochemical Techniques for the Development of
Amperometric Biosensors", Biosensors, 1:161-178 (1985). cited by
other .
Degani, Y. et al., "Direct Electrical Communication between
Chemically Modified Enzymes and Metal Electrodes. 1. Electron
Transfer from Glucose Oxidase to Metal Electrodes via Electron
Relays, Bound Covalently to the Enzyme," J. Phys. Chem.,
91(6):1285-1289 (1987). cited by other .
Degani, Y. et al., "Direct Electrical Communication between
Chemically Modified Enzymes and Metal Electrodes. 2. Methods for
Bonding Electron-Transfer Relays to Glucose Oxidase and
D-Amino-Acid Oxidase," J. Am. Chem. Soc., 110(8):2615-2620 (1988).
cited by other .
Degani, Y. et al., "Electrical Communication between Redox Centers
of Glucose Oxidase and Electrodes via Electrostatically and
Covalently Bound Redox Polymers," J. Am. Chem. Soc., 111:2357-2358
(1989). cited by other .
Denisevich, P. et al., "Unidirectional Current Flow and Charge
State Trapping at Redox Polymer Interfaces on Bilayer Electrodes:
Principles, Experimental Demonstration, and Theory," J. Am. Chem.
Soc., 103(16):4727-4737 (1981). cited by other .
Dicks, J. M., "Ferrocene modified polypyrrole with immobilised
glucose oxidase and its application in amperometric glucose
microbiosensors," Ann. Biol. clin., 47:607-619 (1989). cited by
other .
Ellis, C. D., "Selectivity and Directed Charge Transfer through an
Electroactive Metallopolymer Film," J. Am. Chem. Soc.,
103(25):7480-7483 (1981). cited by other .
Engstrom, R.C., "Electrochemical Pretreatment of Glassy Carbon
Electrodes", Anal. Chem., 54(13):2310-2314 (Nov. 1982). cited by
other .
Engstrom, R.C. et al., "Characterization of Electrochemically
Pretreated Glassy Carbon Electrodes", Anal. Chem., 56(2):136-141
(Feb. 1984). cited by other .
Enthone Inc., "ENPLATE DSR-3241 Cost and Process Control:
Application Process," Imaging Technologies Update. Jun. 2001, No.
3. cited by other .
Feldman, B.J. et al., "Electron Transfer Kinetics at Redox
Polymer/Solution Interfaces Using Microelectrodes and Twin
Electrode Thin Layer Cells", J. Electroanal. Chem., 194(1):63-81
(Oct. 10, 1985). cited by other .
Fischer, H. et al., "Intramolecular Electron Transfer Mediated by
4,4'-Bipyridine and Related Bridging Groups", J. Am. Chem. Soc.,
98(18):5512-5517 (Sep. 1, 1976). cited by other .
Foulds, N.C. et al., "Enzyme Entrapment in Electrically Conducting
Polymers," J. Chem. Soc., Faraday Trans 1., 82:1259-1264 (1986).
cited by other .
Foulds, N.C. et al., "Immobilization of Glucose Oxidase in
Ferrocene-Modified Pyrrole Polymers," Anal. Chem.,
609(22):2473-2478 (Nov. 15, 1988). cited by other .
Frew, J.E. et al., "Electron-Transfer Biosensors", Phil. Trans. R.
Soc. Lond., B316:95-106 (1987). cited by other .
Gamache et al. ("Simultaneous measurement of monamines, metabolites
and amino acids in brain tissue and microdialysis perfusates", J.
Chromatogr., Biomed. Appl. (1993), 614(2), 213-20), 1993. cited by
other .
Gorton, L. et al., "Selective detection in flow analysis based on
the combination of immobilized enzymes and chemically modified
electrodes," Analytica Chimica Acta., 250:203-248 (1991). cited by
other .
Gregg, B. A. et al., "Cross-Linked Redox Gels Containing Glucose
Oxidase for Amperometric Biosensor Applications," Analytical
Chemistry, 62(3):258-263 (Feb. 1, 1990). cited by other .
Gregg, B. A. et al., "Redox Polymer Films Containing Enzymes. 1. A
Redox-Conducting Epoxy Cement: Sythesis, Characterization, and
Electrocatalytic Oxidation of Hydroquinone," J. Phys. Chem.,
95(15):5970-5975 (1991). cited by other .
Groom et al. ("Electrical communication between a water-soluble
1,1'-dimethylferrocene-2-hydroxypropyl-b-cyclodextrin complex and
glucos oxidase; biosensor applications", Biosensors &
Bioelectronics 9:305-313 (1994). cited by other .
Grubb et al., "Blood oxygen content in microliter samples using an
easy-to-build galvanic cell", Journal of Applied Physiology, pp.
456-464 (1981). cited by other .
Hale, P.D. et al., "A New Class of Amperometric Biosensor
Incorporating a Polymeric Electron-Transfer Mediator," J. Am. Chem.
Soc., 111(9):3482-3484 (1989). cited by other .
Harrison, D.J. et al., "Characterization of Perfluorosulfonic Acid
Polymer Coated Enzyme Electrodes and a Miniaturized Integrated
Potentiostat for Glucose Analysis in Whole Blood", Anal. Chem.,
60(19):2002-2007 (Oct. 1, 1988). cited by other .
Hawkridge, F. M. et al., "Indirect Coulometric Titration of
Biological Electron Transport Components," Analytical Chemistry,
45(7):1021-1027 (Jun. 1973). cited by other .
Heineman, W.R. "Spectro-electro-chemistry", Analytical Chemistry,
50(3):390-392, 394, 396, 398, 400, 402 (Mar. 1978). cited by other
.
Heineman, W.R. et al., "Measurement of Enzyme E.degree. Values by
Optically Transparent Thin Layer Electrochemical Cells", Analytical
Chemistry, 47(1):79, 82-84 (Jan. 1975). cited by other .
Heller, A., "Electrical Wiring of Redox Enzymes," Acc. Chem. Res.,
23(5):129-134 (1990). cited by other .
Heller, A., "Electrical Connection of Enzyme Redox Centers to
Electrodes," J. Phys. Chem 96(9):3579-3587 (1992). cited by other
.
Heller, A., "Amperometric biosensors based on three-dimensional
hydrogel-forming epoxy networks," Sensors and Actuators B,
13-14:180-183 (1993). cited by other .
Hubbard, A. et al. "The Theory and Practice of Electrochemistry
with Thin Layer Cells", Electroanalytical Chemistry A Series of
Advances, vol. 4, pp. 129-131, 142-147, 168-171, edited by Allen J.
Bard, Marcel Deckker, Inc. New York (1970). cited by other .
Hubbard, A. et al., "Electrochemistry in Thin Layers of Solution",
CRC Critical Reviews in Analytical Chemistry, 3(2):201-242 (Mar.
1973). cited by other .
Ianniello, R.M. et al., "Differential Pulse Voltammetric Study of
Direct Electron Transfer in Glucose Oxidase Chemically Modified
Graphite Electrodes", Anal. Chem., 54:(7):1098-1101 (Jun. 1982).
cited by other .
Ianniello, R.M. et al. "Immobilized Enzyme Chemically Modified
Electrode as an Amperometric Sensor", Anal. Chem., 53(13):2090-2095
(Nov. 1981). cited by other .
Ikeda, T. et al., "Glucose oxidase-immobilized benzoquinone-carbon
paste electrode as a glucose sensor," Agric. Biol. Chem., 49(2) (1
page--Abstract only) (1985). cited by other .
Ikeda, T. et al., "Kinetics of Outer-Sphere Electon Transfers
Between Metal Complexes in Solutions and Polymeric Films on
Modified Electrodes", J. Am. Chem. Soc., 103(25):7422-7425 (Dec.
16, 1981). cited by other .
Johnson, J. M. et al., "Potential-Dependent Enzymatic Activity in a
Enzyme Thin-Layer Cell," Anal. Chem. 54:1377-1383 (1982). cited by
other .
Johnson, K.W., "Reproducible Electrodeposition of Biomolecules for
the Fabrication of Miniature Electroenzymatic Biosensors", Sensors
and Actuators B Chemical, B5:85-89 (1991). cited by other .
Jonsson, G. et al., "An Amperometric Glucose Sensor Made by
Modification of a Graphite Electrode Surface With Immobilized
Glucose Oxidase and Adsorbed Mediator", Biosensors, 1:35-368
(1985). cited by other .
Josowicz, M. et al., "Electrochemical Pretreatment of Thin Film
Platinum Electrodes", J. Elecrochem. Soc., 135(1):112-115 (Jan.
1988). cited by other .
Karube et al., "Microbiosensors Prepared by Micromachining," GBF
Monographs (1992), 17 (Biosens.: Fundam. Technol. appl.) 477-89.
cited by other .
Katakis, I. et al., "L-.alpha.-Glycerophosphate and L-Lactate
Electrodes Based on the Electrochemical "Wiring" of Oxidases,"
Analytical Chemistry, 64(9):1008-1013 (May 1, 1992). cited by other
.
Katakis, I. et al., "Electrostatic Control of the Electron Transfer
Enabling Binding of Recombinant Glucose Oxidase and Redox
Polyelectrolytes," J. Am. Chem. Soc., 116(8):3617-3618 (1994).
cited by other .
Kenausis, G. et al., "`Wiring` of glucose oxidase and lactate
oxidase within a hydrogel made with poly(vinyl pyridine) complexed
with [Os(4,4'-dimethoxy-2,2'-bipyridine).sub.2 C1].sup.+/2+," J.
Chem. Soc., Faraday Trans., 92(20):4131-4136 (1996). cited by other
.
Kishimoto et al., "Home Care Disposable Glucose Sensor for
Blood-Sugar Monitoring," Sumimoto Met., 46(4) (1994). cited by
other .
Kissinger, "Biomedical Applications of Liquid
Chromatography-Electrochemistry" Journal of Chromatography, 488
(1989) 31-52, month unknown. cited by other .
Koudelka, M. et al., "In-Vivo Behaviour of Hypodermically Implanted
Microfabricated Glucose Sensors", Biosensors & Bioelectronics,
6(1):31-36 (1991). cited by other .
Kulys, J. et al., "Mediatorless peroxidase electode and preparation
of bienzyme sensors," Bioelectrochemistry and Bioenergetics,
24:305-311 (1990). cited by other .
Lager, W. et al., "Implantable Electrocatalytic Glucose Sensor,"
Horm. Metab. Res., 26:526-530 (Nov. 1994). cited by other .
Linder, E. et al., "Flexible (Kapton-Based) Microsensor Arrays of
High Stability for Cardiovascular Applications", J. Chem.
Soc.Faraday Trans., 89(2):361-367 (Jan. 21, 1993). cited by other
.
Maidan, R. et al., "Elimination of Electrooxidizable
Interferant-Produced Currents in Amperometric Biosensros,"
Analytical Chemistry, 64(23):2889-2896 (Dec. 1, 1992). cited by
other .
McNeil, C. J. et al., "Thermostable Reduced Nicotinamide Adenine
Dinucleotide Oxidase: Application to Amperometric Enzyme Assay,"
Anal. Chem., 61(1):25-29 (Jan. 1, 1989). cited by other .
Miyasaka, Takehiro, "Development of Enzyme Controlled Glucose
Sensor in Blood Activated . . . ," Chemical Engineering, vol. 42,
No. 5 (Jun. 19, 1995). cited by other .
Miyawaki, O. et al., "Electrochemical and Glucose Oxidase Coenzyme
Activity of Flavin Adenine Dinucleotide Covalently Attached to
Glassy Carbon at the Adenine Amino Group", Biochimica et Biophysica
Acta, 838:60-68 (1985). cited by other .
Moatti-Sirat, D. et al., "Evaluating in vitro and in vivo the
inteference of ascorbate and acetaminophen on glucose detection by
a needle-type glucose sensor," Biosensors & Bioelectronics,
7(5):345-352 (1992). cited by other .
Moatti-Sirat, D. et al., "Towards continuous glucose monitoring: in
vivo evaluation of a miniaturized glucose sensor implanted for
several days in rat subcutaneous tissue," Diabetologia, 35(3) (1
page--Abstract only) (Mar. 1992). cited by other .
Moatti-Sirat, D. et al., "Reduction of acetaminophen interference
in glucose sensors by a composite Nafion membrane: demonstration in
rats and man," Diabetologia, 37(6) (1 page--Abstract only) (Jun.
1994). cited by other .
Morris, N.A., "An Electrochemical Capillary Fill Device for the
Analysis of Glucose Incorporating Glucose Oxidase and Ruthenium
(III) Hexamine as Mediator," Electroanalysis, vol. 4, No. 1 (Jan
1992). cited by other .
Moser, I. et al., "Advanced Immobilization and Protein Techniques
on thin Film Biosensors", Sensors and Actuators, B7:356-362 (1992).
cited by other .
Moser et al., "Rapid liver enzyme assay with miniaturized liquid
handling system comprising thin film biosensor array", Sensors and
Actuators B: Chemical 1997;44(1-3):377-80. cited by other .
Moussy, F. et al., "Performance of Subcutaneously Implanted
Needle-Type Glucose Sensors Employing a Novel Trilayer Coating",
Anal. Chem., 65:2072-2077 (1993). cited by other .
Nagy, G. et al., "A New Type of Enzyme Electrode: The Ascorbic Acid
Eliminator Electrode," Life Sciences, 31(23):2611-2616 (1982).
cited by other .
Nakamura, S. et al., "Effect of Periodate Oxidation on the
Structure and Properties of Glucose Oxidase," Biochimica et
Biophysica Acta., 445:294-308 (1976). cited by other .
Narazimhan, K. et al., "p-Benzoquinone activation of metal oxide
electrodes for attachment of enzymes," Enzyme Microb. Technol.,
7(6) (1 page--Abstract only) (1985). cited by other .
Niwa et al. "Small-Volume Voltammetric Detection of 4-aminophenol
with Interdigitated Array Electrodes and its Application to
Electrochemcial Enzyme Immunoassay," Anal. Chem., (Jun. 1993), 65,
1559-1563. cited by other .
Niwa, O. et al., "Highly Sensitive Small Volume Voltammetry of
Reversible Redox Species with and IDA Electrochemical Cell and its
Application to Selective Detection of Catecholamine", Sensors and
Actuators B, 13-14, pp. 558-560 (1993). cited by other .
Niwa, O. et al., "Concentration of Extracellular L-Glutamate
Released from Cultured Nerve Cells Measured with a Small-Volume
Online Sensor," Analytical Chemistry, 68(11), Jun. 1, 1996, pp.
1865-1870. cited by other .
Ohara, T. J. et al., ""Wired" Enzyme Electrodes for Amperometric
Determination of Glucose or Lactate in the Presence of Interfering
Substances," Analytical Chemistry, 66(15):2451-2457 (Aug. 1, 1994).
cited by other .
Ohara, T. J., "Osmium Bipyridyl Redox Polymers Used in Enzyme
Electrodes," Platinum Metals Rev., 39(2):54-62 (Apr. 1995). cited
by other .
Olievier, C. N. et al., "In vivo Measurement of Carbon Dioxide
Tension with a Miniature Electrode," Pflugers Arch. 373:269-272
(1978). cited by other .
Paddock, R. et al., "Electrocatalytic reduction of hydrogen
peroxide via direct electron transfer from pyrolytic graphite
electrodes to irreversibly adsorbed cytochrome c peroxidase," J.
Electroanal. Chem., 260:487-494 (1989). cited by other .
Palleschi, G. et al., "A Study of Interferences in Glucose
Measurements in Blood by Hydrogen Peroxide Based Glucose Probes",
Anal. Biochem., 159:114-121 (1986). cited by other .
Pickup, J. et al., "Potentially-implantable amperometric glucose
sensors with mediated electron transfer: improving the operating
stability," Biosensors, 4(2) (1 page--Abstract only) (1989). cited
by other .
Pickup, J. C. et al., "In vivo molecular sensing in diabetes
mellitus: an implantable glucose sensor with direct electron
transfer," Diabetologia, 32(3):213-217 (1989). cited by other .
Pickup, J., "Developing glucose sensors for in vivo use," TIBTECH,
11(7): 285-289 (Jul. 1993). cited by other .
Pishko, M.V. et al., "Amperometric Glucose Microelectrodes Prepared
Through Immobilization of Glucose Oxidase in Redox Hydrogels",
Anal. Chem., 63(20):2268-2272 (Oct. 15, 1991). cited by other .
Poitout, V. et al., "In vitro and in vivo evaluation in dogs of a
miniaturized glucose sensor," ASAIO Transactions, 37(3) (1
page--Abstract only) (Jul.-Sep. 1991). cited by other .
Poitout, V. et al., "Calibration in dogs of a subcutaneous
miniaturized glucose sensor using a glucose meter for blood glucose
determination," Biosensors & Bioelectronics, 7:587-592 (1992).
cited by other .
Poitout, V. et al., "A glucose monitoring system for on line
estimation in man of blood glucose concentration using a
miniaturized glucose sensor implanted in the subcutaneous tissue
and a wearable control unit," Diabetolgia, 36(7) (1 page--Abstract
only) (Jul. 1993). cited by other .
Pollak, A. et al., "Enzyme Immobilization by Condensation
Copolymerization into Cross-Linked Polyacrylamide Gels," J. Am.
Chem. Soc., 102(20):6324-6336 (1980). cited by other .
Pons, B. S. et al., "Application of Deposited Thin Metal Films as
Optically Transparent Electrodes for Internal Reflection
Spectometric Observation of Electrode Solution Interfaces",
Analytical Chemistry, 39(6):685-688, (May 1967). cited by other
.
Reach, G. et al., "A Method for Evaluating in vivo the Functional
Characteristics of Glucose Sensors", Biosensors 2:211-220 (1986).
cited by other .
Reach, G. et al., "Can Continuous Glucose Monitoring Be Used for
the Treatment of Diabetes?" Analytical Chemistry, 64(6):381-386
(Mar. 15, 1992). cited by other .
Rebrin, K. et al., "Automated Feedback Control of Subcutaneous
Glucose Concentration in Diabetic Dogs", Diabetologia,
32(8):573-576 (Aug. 1989). cited by other .
Reilley, "Electrochemistry Using Thin-Layer Cells", Rev. Pure and
Appl. Chem., 18, pp. 137-151 (1968). cited by other .
Roche's Final Invalidity Contentions of '745 and '551 Patents as of
Jun. 18, 2007, and references. cited by other .
Roe, "Comparison of Amperometric and Coulometric Electrochemical
Detectors for HPLC through a Figure of Merit", Analytical Letters,
16(A8), 613-631 (1983). cited by other .
Sakakida, M. et al., "Ferrocene-mediate needle-type glucose sensor
covered with newly designed biocompatible membrane," Sensors and
Actuators .beta., 13-14:319-322 (1993). cited by other .
Samuels, G. J. et al., "An Electrode-Supported Oxidation Catalyst
Based on Ruthenium (IV). pH "Encapsulation" in a Polymer Film," J.
Am. Chem. Soc., 103(2):307-312 (1981). cited by other .
Sasso, S.V. et al., "Electropolymerized 1,2-Diaminobenzene as a
Means to Prevent Interferences and Fouling and to Stabilize
Immobilized Enzyme in Electrochemical Biosensors", Anal. Chem.,
62(11): 1111-1117 (Jun. 1, 1990). cited by other .
Schmehl, R.H. et al., "The Effect of Redox Site Concentration on
the Rate of Mediated Oxidation of Solution Substrates by a Redox
Copolymer Film", J. Electroanal. Chem., 152:97-109 (Aug. 25, 1983).
cited by other .
Shichiri, M. et al., "Glycemic Control in Pancreatectomized Dogs
with a Wearable Artificial Endocrine Pancreas", Diabetologia,
24(3):179-184 (Mar. 1983). cited by other .
Shigeru, T. et al, "Simultaneous Determination of Glucse and
1,5-=Anydroglucitol", Chemical Abstracts, 111:394 (1989). cited by
other .
Sittampalam, G. et al., "Surface-Modified Electrochemical Detector
for Liquid Chromatography", Anal. Chem., 55(9):1608-1610 (Aug.
1983). cited by other .
Sprules, S. D. et al., "Evaluation of a New Disposable
Screen-Printed Sensor Strip for the Measurement of NADH and Its
Modification to Produce a Lactate Biosensor Employing Microliter
Volumes," Electroanalysis, 8(6):539-543 (1996). cited by other
.
Sternberg, R. et al., "Study and Development of Multilayer
Needle-type Enzyme-based Glucose Microsensors," Biosensors, 4:27-40
(1988). cited by other .
Sternberg, R. et al., "Covalent Enzyme Coupling on Cellulose
Acetate Membranes for Glucose Sensor Development," Analytical
Chemistry, 60(24):2781-2786 (Dec. 15, 1988). cited by other .
Sternberg, F. et al., "Calibration Problems of Subcutaneous
Glucosensors when Applied "In-Situ" in Man," Norm. metabl. Res,
26:523-525 (1994). cited by other .
Suekane, M., "Immobilization of glucose isomerase," Zeitschrift fur
Allgemeine Mikrobiologie, 22(8):565-576 (1982). cited by other
.
Tajima, S. et al., "Simultaneous Determination of Glucose and
1.5-Anydrogluc.sub.--tol", Chemical Abstracts, 111(25):394
111:228556g (Dec. 18, 1989). cited by other .
Takata, Y., "Liquid Chromatography with Coulometric Detector",
Advances in Liquid Chromatography: 35 years of col. Liquid, Eds.
Hanai et al., World Scientific, pp. 43-74 (1996). cited by other
.
Tarasevich, M.R. "Bioelectrocatalysis", Comprehensive Treatise of
Electrochemistry, 10 (Ch. 4):231-295 (1985). cited by other .
Tatsuma, T. et al., "Enzyme Monolayer- and Bilayer-Modified Tin
Oxide Electrodes for the Determination of Hydrogen Peroxide and
Glucose," Anal. Chem., 61(2):2352-2355 (Nov. 1, 1989). cited by
other .
Taylor, C. et al., "`Wiring` of glucose oxidase within a hydrogel
made with polyvinyl imidazole complexed with
[(Os-4,4'-dimethoxy-2,2'-bipyridine)C1].sup.+/2+," Journal of
Electroanalytical Chemistry, 396:511-515 (1995). cited by other
.
Therasense, Inc. (now known as Abbott Diabetes Care Inc.) and
Abbott Laboratories v. Becton, Dickinson and Company and Nova
Biomedical Corporation, and Bayer Healthcare LLC, United States
Court of Appeals for the Federal Circuit, Decided: Jan. 25, 2010.
cited by other .
Tietz, in: "Textbook of Clinical Chemistry", C. A. Burtis and E.R.
Ashwood, eds., W. B. Saunders Co., Phila 1994, pp. 2210-2212. cited
by other .
Trojanowicz, M. et al., "Enzyme Entrapped Polypyrrole Modified
Electrode for Flow-Injection Determination of Glucose," Biosensors
& Bioelectronics, 5:149-156 (1990). cited by other .
Turner, "Research: A new approach to blood glucose tests", Balance,
(Aug. 1983). cited by other .
Turner, A.P.F. et al., "Diabetes Mellitus: Biosensors for Research
and Management", Biosensors, 1:85-115 (1985). cited by other .
Turner, R. F. B. et al., "A Biocompatible Enzyme Electrode for
Continuous in vivo Glucose Monitoring in Whole Blood," Sensors and
Actuators, B1(1-6):561-564 (Jan. 1990). cited by other .
Tuzhi, P. et al., "Constant Potential Pretreatment of Carbon Fiber
Electrodes for In Vivo Electrochemistry", Analytical Letters,
24(6):935-945 (1991). cited by other .
Uhegbu et al., "Initial Studies of a New Approach to the Design and
Use of Enzyme-Based Reactor/Sensor Systems: Amperometric System for
Glucose" Anal. Chem. 1993, 65, 2443-2451, month unknown. cited by
other .
Umaha, M., "Protein-Modified Electrochemically Active Biomaterial
Surface," U.S. Army Research Office Report, (12 pages) (Dec. 1988).
cited by other .
Urban, G. et al., "Miniaturized Thin-Film Biosensors Using
Covalently Immobilized Glucose Oxidase", Biosensors &
Bioelectronics, 6(7):555-562 (1991). cited by other .
Van Der Schoot et al., "An ISFET-Based Microlite Titrator:
Integration of a Chemical Sensor-Actuator System", Sensors and
Actuators, 8:11-22 (1985). cited by other .
Velho, G. et al., "In Vitro and In Vivo Stability of Electrode
Potentials in Needle-Type Glucose Sensors", Diabetes, 38(2):164-171
(Feb. 1989). cited by other .
Velho, G. et al., "Strategies for calibrating a subcutaneous
glucose sensor," Biomed. Biochin. Acta, 48(11/12):957-964 (1989).
cited by other .
Vidal, J.C. et al., "A chronoamperometric sensor for hydrogen
peroxide based on electron transfer between immobilized horseradish
peroxidase on a glassy carbon electrode and a diffusing ferrocene
mediator", Sensors and Actuators B 21, pp. 135-141 (1994). cited by
other .
Von Woedtke, T. et al., "In Situ Calibration of Implanted
Electrochemical Glucose Sensors," Biomed. Biochim. Acta,
48(11/12):943-952 (1989). cited by other .
Vreeke, M. et al., "Hydrogen Peroxide and .beta.-Nicotinamide
Adenine Dinucleotide Sensing Amperometric Electrodes Based on
Electrical Connection of Horseradish Peroxidase Redox Centers to
Electrodes through a Three-Dimensional Electron Relaying Polymer
Network," Analytical Chemistry, 64(24):3084-3090 (Dec. 15, 1992).
cited by other .
Vreeke, M. S. et al., "Chapter 15: Hydrogen Peroxide Electrodes
Based on Electrical Connection of Redox Centers of Various
Peroxidases to Electrodes through a Three-Dimensional Electron
Relaying Polymer Network," Diagnostic Biosensor Polymers, 7 pp.
(Jul. 26, 1993). cited by other .
Vriend, J., "Determination of Amino Acids and Monoamine
Neurotransmitters in Caudate Nucleus of Seizure-Resistant and
Seizure-Prone BALB/c Mice," Journal of Neurochemistry, vol. 60, No.
4, pp. 1300-1307 (1993). cited by other .
Wang, J. et al., "Activation of Glassy Carbon Electrodes by
Alternating Current Electrochemical Treatment", Analytica Chimica
Acta, 167:325-334 (Jan. 1985). cited by other .
Wang, J. et al., "Amperometric biosensing of organic peroxides with
peroxidase-modified electrodes," Analytica Chimica Acta. 254:81-88
(1991). cited by other .
Wang, J. et al., "Screen-Printable Sol-Gel Enzyme-Containing Carbon
Inks," Analytical Chemistry, 68(15):2705-2708 (Aug. 1, 1996). cited
by other .
Wang, J. et al., "Sol-Gel-Derived Metal-Dispersed Carbon Composite
Amperometric Biosensors," Electroanalysis, 9(1):52-55 (1997). cited
by other .
Williams, D.L. et al., "Electrochemical-Enzymatic Analysis of Blood
Glucose and Lactate", Anal. Chem., 42(1):118-121 (Jan. 1970). cited
by other .
Wilson, G. S. et al., "Progress toward the Development of an
Implantable Sensor for Glucose," Clinical Chemistry,
38(9):1613-1617 (1992). cited by other .
Yabuki, S. et al., "Electro-conductive Enzyme Membrane," J. Chem.
Soc. Chem. Commun, 945-946 (1989). cited by other .
Yang et al., "Application of "Wired" Peroxidase Electrodes for
Peroxide Determination in Liquid Chromatography Coupled to Oxidase
Immobilized Enzyme Reactors" Anal. Chem. 1995, 67, 1326-1331 Apr.
cited by other .
Yang et al., "Continuous Monitoring of Subcutaneous Glucose and
Lactate Using Microdialysis With On-Line Enzyme Electrodes" Current
Separations 14:1(1995) pp. 31-35, month unknown. cited by other
.
Yang, L. et al., "Determination of Oxidase Enzyme Substrates Using
Cross-Flow Thin-Layer Amperometry," Electroanalysis, 8(8-9):716-721
(1996). cited by other .
Yao, S.J. et al., "The Interference of Ascorbate and Urea in
Low-Potential Electrochemical Glucose Sensing", Proceedings of the
Twelfth Annual International Conference of the IEEE Engineering in
Medicine and Biology Society, 12(2):487-489 (Nov. 1-4, 1990). cited
by other .
Ye, L. et al., "High Current Density "Wired" Quinoprotein Glucose
Dehydrogenase Electrode," Anal. Chem., 65(3):238-241 (Feb. 1,
1993). cited by other .
Yildiz, A., "Evaluation of an Improved Thin-Layer Electrode",
Analytical Chemistry, 40(7):1018-1024 (Jun. 1968). cited by other
.
Zamzow, K. et al., New Wearable Continuous Blood Glucose Monitor
(BGM) and Artificial Pancreas (AP), Diabetes, 39:5A(20) (May 1990).
cited by other .
Zhang, Y. et al. "Application of cell culture toxicity tests to the
development of implantable biosensors," Biosensors &
Bioelectronics, 6:653-661(1991). cited by other .
Batchelor et al., "Amperometric Assay for the Ketone body
3-Hydroxybutyrate," Analytica Chimica Acta, 221:289-294 (1989).
cited by other .
Choleau et al., "Calibration of a subcutaneous amperometric glucose
sensor Part 1. Effect of measurement uncertainties on the
determination of sensor sensitivity and background current,"
Biosensors and Bioelectronics, 17:641-646 (2002). cited by other
.
Kerner et al., "The function of a hydrogen peroxide-detecting
electroenzymatic glucose electrode is markedly impaired in human
sub-cutaneous tissue and plasma," Biosensors & Bioelectronics,
8:473-482 (1993). cited by other .
Linke et al., "Prevention of the Decrease in Sensitivity of an
Amperometric Glucose Sensor in Undiluted Human Serum," Clinical
Chemistry, 45( 2):283-285 (1999). cited by other .
Pickup et al., "Responses and calibration of amperometric glucose
sensors implanted in the subcutaneous tissue of man," Acta.
Diabetol., 30:143-148 (1993). cited by other .
Pravda, M. et al., "Evaluation of amperometric glucose biosensors
based on co-immobilisation of glucose oxidase with an osmium redox
polymer in electrochemically generated polyphenol films," Analytica
Chimica Acta, 304:127-138 (1995). cited by other .
TheraSense, Inc., v. Becton, Dickinson and Co., 560 F. Supp. 2d
835--Dist. Court, ND California (Apr. 3, 2008). cited by other
.
Thome-Duret et al., "Modification of the Sensitivity of Glucose
Sensor Implanted into Subcutaneous Tissue," Diabetes &
Metabolism (Paris), 22:174-1 78 (1996). cited by other .
Ward et al., "A new amperometric glucose microsensor: in vitro and
short-term in vivo evaluation," Biosensors & Bioelectronics,
17:181-189 (2002). cited by other .
Abruna, H. D. et al., "Rectifying Interfaces Using Two-Layer Films
of Electrochemically Polymerized Vinylpyridine and Vinylbipyridine
Complexes of Ruthenium and Iron on Electrodes," J. Am. Chem. Soc.,
103(1):1-5 (Jan. 14, 1981). cited by other .
Albery, W. J. et al., "Amperometric enzyme electrodes. Part II.
Conducting salts as electrode materials for the oxidation of
glucose oxidase," J. Electroanal. Chem. Interfacial Electrochem.,
194(2) pp. 223-235 (1985). cited by other .
Albery, W. J. et al., "Amperometric Enzyme Electrodes," Phil.
Trans. R. Soc. Lond. B316:107-119 (1987). cited by other .
Alcock, S. J. et al., "Continuous Analyte Monitoring to Aid
Clinical Practice," IEEE Engineering in Medicine and Biology,
319-325 (1994). cited by other .
Anderson, L. B. et al., "Thin-Layer Electrochemistry: Steady-State
Methods of Studying Rate Processes," J. Electroanal. Chem.,
10:295-305 (1965). cited by other .
Anderson, C. W. et al., "A Small-Volume Thin-Layer
Spectroelectrochemical Cell for the Study of Biological
Components", Analytical Biochemistry, 93(2):366-372 (1979). cited
by other .
Bard and Faulkner, "Electrochemical Methods: Fundamentals and
Applications", pp. 2-3, 23-24 (1980). cited by other .
Bartlett, P. N. et al., "Covalent Binding of Electron Relays to
Glucose Oxidation," J. Chem. Soc. Chem. Commun., 1603-1604 (1987).
cited by other .
Bartlett, P. N. et al., "Strategies for the Development of
Amperometric Enzyme Electrodes," Biosensors, 3:359-379 (1987/1988).
cited by other .
Bartlett, P. N. et al., "Modification of glucose oxidase by
tetrathiafulvalene," J. Chem. Soc., Chem. Commun., 16:1135-1136
(1990). cited by other .
Bayer Corporation, Glucometer DEX blood glucose monitoring system,
User guide. Bayer Corporation, (Rev. Jul. 1997). cited by other
.
Bayer Corporation, Glucometer elite diabetes care system, User
guide for use with Glucometer Elite blood glucose meter. Bayer
Corporation, (Rev. Jun. 1998). cited by other .
Bayer's Invalidity Contentions of '745 and '551 Patents as of Jun.
18, 2007, and references. cited by other .
Caglar and Wnek, "Glucose-Sensitive Polyphyrrole/poly
(Styrenesulfonate) Films Containing Co-Immobilized Glucose Oxidase
and (Ferrocenylmethyl) Trimethylammonium Bromide," J. of
Macromolecular Sc.--Pure Appl. Chem., A32(2), pp. 349-359 (1995).
cited by other .
Chen, C.Y. et al., "Amperometric Needle-Type Glucose Sensor based
on a Modified Platinum Electrode with Diminished Response to
Interfering Materials", Analytica Chimica Acta, 265:5-14 (1992).
cited by other .
Clark, L.C., Jr. et al., "Electrode Systems for Continuous
Monitoring in Cardiovascular Surgery," Annals New York Academy of
Sciences, pp. 29-45 (1962). cited by other .
Diffusion coefficient of glucose in water--Homo sapiens from the
Catalog of useful Biological Numbers at harvard.edu downloaded from
http://bionumbers.hms.harvard.edu/ on Nov. 30, 2011. cited by other
.
Gernet, S. et al., "Fabrication and Characterization of a Planar
Electrochemical Cell and Its Application as a Glucose Sensor",
Biosensors & Actuators, 18:59-70 (1989). cited by other .
Huang et al., "Detection of basal acetylcholine in rat brain
microdialyse", Journal of Chromatography B, 670. 323-327 (1995).
cited by other .
Jobst et al., Mass producible miniaturized flow through a device
with a biosensor array, Sensors and Actuators B: 43 (Sep. 1997)
121-125. cited by other .
Johnson K. W. et al., "In Vivo Evaluation of an Electroenzymatic
Glucose Sensor Implanted in Subcutaneous Tissue", Biosensors &
bioelectronics 7:709-714 (1992). cited by other .
Kondo, T. et al., "A Miniature Glucose Sensor, Implantable in the
Blood Stream", Diabetes Care, 5(3):218-221 (May-Jun. 1982). cited
by other .
Koudelka, M. et al., "In-Vivo Behaviour of Hypodermically Implanted
Microfabricated Glucose Sensors", Biosensors & Bioelectronics,
24:305-311 (1990). cited by other .
Kuhn L. S., "Biosensors: blockbuster or bomb? Electrochemical
biosensors for diabetes monitoring," The Electrochemical Society
Interface, 26-31 (1998). cited by other .
Lee, J. et al., "A New Glucose Sensor using Microporous Enzyme
Membrane", Sensors and Actuators, B3:215-219 (1991). cited by other
.
Lewandowski, J.J. et al., "Evaluation of a Miniature Blood Glucose
Sensor", Trans Am Soc Artif Intern Organs, XXXIV: 255-258 (1988).
cited by other .
Liu et al., "Miniature Multiple Cathode Dissolved Oxygen Sensor for
Marine Science Applications", Marine Technology "The Decade of
Oceans" pp. 468-472 (1980). cited by other .
Liu and Neuman, "Fabrication of Miniature PO2 and pH Sensors Using
Microelectronic Techniques", Diabetes Care, vol. 5, No. 3, pp.
275-276 (May-Jun. 1982). cited by other .
Mann-Buxbaum, E. et al, "New Microminiaturized Glucose Sensors
Using Covalent Immobilization Techniques", Sensors and Actuators,
B1:518-522 (1990). cited by other .
Mastrototaro, J.J. et al., "An Electroenzymatic Glucose Sensor
Fabricated on a Flexible Substrate", Sensors and Biosensors B
Chemical, B5:139-144 (1991). cited by other .
Matthews, D.R., et al., "An Amperometric Needle-Type Glucose Sensor
Tested in Rats and Man", Original Articles, pp. 248-252 (1988).
cited by other .
McDuffie et al., "Twin Electrode Thin Layer Electrochemistry:
Determination of Chemical Reaction Rates by Decay of Steady-State
Current", Analytical Chemistry, vol. 38, No. 7, pp. 883-890 (Jun.
1966). cited by other .
McKean et al., "A telemetry-Instrumentation System for Chronically
Implanted Glucose and Oxygen Sensors", IEEE Transactions of
Biomedical Engineering, 35(7):526-532 (Jul. 1988). cited by other
.
Ohara, T. J. et al., "Glucose Electrodes Based on Cross-Linked
[Os(bpy).sub.2 Cl].sup.+/2+ Complexed Poly(1-vinylimadazole)
Films," Analytical Chemistry, 65(23):3512-3516 (Dec. 1, 1993).
cited by other .
Palleschi, G. et al., "Ideal Hydrogen Peroxide-Based Glucose
Sensor", Applied Biochemistry and Biotechnology, 31:21-35 (1991).
cited by other .
Pankratov, I. et al., "Sol-gel derived renewable-surface
biosensors," Journal of Electroanalytical Chemistry, 393:35-41
(1995). cited by other .
Pathak, C. P. et al., "Rapid Photopolymerization of
Immunoprotective Gels in Contact with Cells and Tissue," J. Am.
Chem. Soc., 114(21):8311-8312 (1992). cited by other .
Schalkhammer, T. et al., "Electrochemical Glucose Sensors on
Permselective Non-conducting Substituted Pyrrole Polymers", Sensors
and Actuators, B4:273-281 (1991). cited by other .
Scheller, F. et al., "Enzyme electrodes and their application,"
Phil. Trans. R. Soc. Lond., B 316:85-94 (1987). cited by other
.
Soegijoko, S. et al., Horm. Metabl. Res., Suppl. Ser, 12, pp.
165-169 (1982). cited by other .
Turner, A.P.F., "Redox Mediators and their Application in
Amperometric Sensors," Proc. NATO Advanced Research Workshop on
Analytical Uses of Immobilized Biological Compounds for Detection,
Medical and Industrial Uses, Florence, Italy, (May 4-8, 1987), Ed.
Guilbault et al., D. Reidel Publishing Company, pp. 131-140. cited
by other .
Unwin et al., "A Generally Applicable Method for the Measurement of
Heterogeneous Rate Constants of Reactions Occurring at the
Solid/Liquid Interface," Journal of Colloid and Interface Science,
128(1):208-222 (Mar. 1989). cited by other .
Wang, D. L. et al., "Miniaturized Flexible Amperometric Lactate
Probe," Analytical Chemistry, 65(8):1069-1073 (Apr. 15, 1993).
cited by other .
Wingard, "Immobilized enzyme electrode for glucose determination
for the artificial pancreas", Federation Proceedings from
symposiums for Drugs and Enzymes Attached to Solid Supports, pp.
288-291 (1983). cited by other .
Woodard and Reilley, Comprehensive Treatise of Electrochemistry,
Chapter 6 "Thin Layer Cell Techniques", pp. 353-392 (1984). cited
by other .
Yamasaki, Y., "The Development of a Needle-Type Glucose Sensor for
Wearable Artificial Endocrine Pancreas", Medical Journal of Osaka
University, vol. 35, No. 1-2, pp. 24-34 (Sep. 1994). cited by other
.
Yao, T. et al., "A Chemically-Modified Enzyme Membrane Electrode as
an Amperometric Glucose Sensor," Analytica Chimica Acta., 148:27-33
(1983). cited by other .
Zhang, Y. et al., "Elimination of the Acetaminophen Interference in
an Implantable Glucose Sensor," Anal. Chem. 66:1183-1188 (1994).
cited by other.
|
Primary Examiner: Noguerola; Alex
Attorney, Agent or Firm: Baba; Edward J. Davy; Brian E.
Bozicevic, Field & Francis LLP
Parent Case Text
This application is a continuation of Ser. No. 11/734,979, filed on
Apr. 13, 2007, which is a divisional of U.S. application Ser. No.
10/662,081, filed Sep. 12, 2003, now issued as U.S. Pat. No.
7,225,535, which is a divisional application of U.S. application
Ser. No. 09/594,285, filed Jun. 15, 2000, now issued as U.S. Pat.
No. 6,618,934, which is a continuation of U.S. application Ser. No.
09/295,962, filed Apr. 21, 1999, now issued as U.S. Pat. No.
6,338,790, which claims the benefit of U.S. Provisional Application
Ser. No. 60/103,627, filed Oct. 8, 1998 and U.S. Provisional
Application Ser. No. 60/105,773, filed Oct. 8, 1998.
Claims
We claim:
1. A method of determining a concentration of an analyte in a
sample using an electrochemical sensor, the method comprising:
contacting a sample with an electrochemical sensor; drawing at
least a portion of the sample through one of a first and a second
aperture and into a sample chamber; observing a signal from a first
indicator electrode to signify that the sample chamber is beginning
to fill with sample; observing a signal from a second indicator
electrode to signify that the sample chamber is sufficiently full
for determining the concentration of the analyte using the sensor;
and generating a analyte-responsive signal from the sensor in
response to electrolysis of the analyte in the sample to determine
the concentration of the analyte in the sample.
2. The method of claim 1, wherein generating a analyte-responsive
signal comprises generating the ketone-responsive signal using the
working electrode, the counter electrode, and at least one of the
first and second indicator electrodes.
3. The method of claim 1, wherein the step of generating a
analyte-responsive signal occurs simultaneously with and continues
after the step of observing the second indicator electrode.
4. The method of claim 1, wherein the step of generating a
analyte-responsive signal occurs after the step of observing the
second indicator electrode.
5. The method of claim 1, wherein the concentration of the analyte
is determined using coulometry.
6. The method of claim 1, wherein the concentration of the analyte
is determined using amperometry.
7. The method of claim 1, wherein the concentration of the analyte
is determined using potentiometry.
8. The method of claim 1, wherein the analyte is glucose or a
ketone.
9. The method of claim 8, wherein the analyte is glucose.
10. The method of claim 9, wherein the sensor comprises a
glucose-responsive enzyme selected from glucose dehydrogenase and
glucose oxidase.
11. The method of claim 1, further comprising obtaining the sample
from a finger of a subject.
12. The method of claim 1, further comprising obtaining the sample
from a region of a subject having a lower nerve end density as
compared to a fingertip.
13. The method of claim 12, wherein the region of a subject having
a lower nerve end density as compared to a fingertip is selected
from the group consisting of: a forearm region, and a thigh region.
Description
FIELD OF THE INVENTION
This invention relates to analytical sensors for the detection of
bioanalytes in a small volume sample.
BACKGROUND OF THE INVENTION
Analytical sensors are useful in chemistry and medicine to
determine the presence and concentration of a biological analyte.
Such sensors are needed, for example, to monitor glucose in
diabetic patients and lactate during critical care events.
Currently available technology measures bioanalytes in relatively
large sample volumes, e.g., generally requiring 3 microliters or
more of blood or other biological fluid. These fluid samples are
obtained from a patient, for example, using a needle and syringe,
or by lancing a portion of the skin such as the fingertip and
"milking" the area to obtain a useful sample volume. These
procedures are inconvenient for the patient, and often painful,
particularly when frequent samples are required. Less painful
methods for obtaining a sample are known such as lancing the arm or
thigh, which have a lower nerve ending density. However, lancing
the body in the preferred regions typically produces submicroliter
samples of blood, because these regions are not heavily supplied
with near-surface capillary vessels.
It would therefore be desirable and very useful to develop a
relatively painless, easy to use blood analyte sensor, capable of
performing an accurate and sensitive analysis of the concentration
of analytes in a small volume of sample.
Sensors capable of electrochemically measuring an analyte in a
sample are known in the art. Some sensors known in the art use at
least two electrodes and may contain a redox mediator to aid in the
electrochemical reaction. However, the use of an electrochemical
sensor for measuring analyte in a small volume introduces error
into the measurements. One type of inaccuracy arises from the use
of a diffusible redox mediator. Because the electrodes are so close
together in a small volume sensor, diffusible redox mediator may
shuttle between the working and counter electrode and add to the
signal measured for analyte. Another source of inaccuracy in a
small volume sensor is the difficulty in determining the volume of
the small sample or in determining whether the sample chamber is
filled. It would therefore be desirable to develop a small volume
electrochemical sensor capable of decreasing the errors that arise
from the size of the sensor and the sample.
SUMMARY OF THE INVENTION
The sensors of the present invention provide a method for the
detection and quantification of an analyte in submicroliter
samples. In general, the invention includes a method and sensor for
analysis of an analyte in a small volume of sample by, for example,
coulometry, amperometry and/or potentiometry. A sensor of the
invention utilizes a non-leachable or diffusible redox mediator.
The sensor also includes a sample chamber to hold the sample in
electrolytic contact with the working electrode. In many instances,
the sensor also contains a non-leachable or diffusible second
electron transfer agent.
In a preferred embodiment, the working electrode faces a counter
electrode, forming a measurement zone within the sample chamber,
between the two electrodes, that is sized to contain no more than
about 1 .mu.L of sample, preferably no more than about 0.5 .mu.L,
more preferably no more than about 0.25 .mu.L, and most preferably
no more than about 0.1 .mu.L, of sample. A sorbent material is
optionally positioned in the sample chamber and measurement zone to
reduce the volume of sample needed to fill the sample chamber and
measurement zone.
In one embodiment of the invention, a biosensor is provided which
combines coulometric electrochemical sensing with a non-leachable
or diffusible redox mediator to accurately and efficiently measure
a bioanalyte in a submicroliter volume of sample. The preferred
sensor includes an electrode, a non-leachable or diffusible redox
mediator on the electrode, a sample chamber for holding the sample
in electrical contact with the electrode and, preferably, sorbent
material disposed within the sample chamber to reduce the volume of
the chamber. The sample chamber, together with any sorbent
material, is sized to provide for analysis of a sample volume that
is typically no more than about 1 .mu.L, preferably no more than
about 0.5 .mu.L, more preferably no more than about 0.25 .mu.L, and
most preferably no more than about 0.1 .mu.L. In some instances,
the sensor also contains a non-leachable or diffusible second
electron transfer agent.
One embodiment of the invention includes a method for determining
the concentration of an analyte in a sample by, first, contacting
the sample with an electrochemical sensor and then determining the
concentration of the analyte. The electrochemical sensor includes a
facing electrode pair with a working electrode and a counter
electrode and a sample chamber, including a measurement zone,
positioned between the two electrodes. The measurement zone is
sized to contain no more than about 1 .mu.L of sample.
The invention also includes an electrochemical sensor with two or
more facing electrode pairs. Each electrode pair has a working
electrode, a counter electrode, and a measurement zone between the
two electrodes, the measurement zone being sized to hold no more
than about 1 .mu.L of sample. In addition, the sensor also includes
a non-leachable redox mediator on the working electrode of at least
one of the electrode pairs or a diffusible redox mediator on a
surface in the sample chamber or in the sample.
One aspect of the invention is a method of determining the
concentration of an analyte in a sample by contacting the sample
with an electrochemical sensor and determining the concentration of
the analyte by coulometry. The electrochemical sensor includes an
electrode pair with a working electrode and a counter electrode.
The sensor also includes a sample chamber for holding a sample in
electrolytic contact with the working electrode. Within the sample
chamber is sorbent material to reduce the volume sample needed to
fill the sample chamber so that the sample chamber is sized to
contain no more than about 1 .mu.L of sample. The sample chamber
also contains a non-leachable or diffusible redox mediator and
optionally contains a non-leachable or diffusible second electron
transfer agent.
The sensors may also include a fill indicator, such as an indicator
electrode or a second electrode pair, that can be used to determine
when the measurement zone or sample chamber has been filled. An
indicator electrode or a second electrode pair may also be used to
increase accuracy of the measurement of analyte concentration. The
sensors may also include a heating element to heat the measurement
zone or sample chamber to increase the rate of oxidation or
reduction of the analyte.
Sensors can be configured for side-filling or tip-filling. In
addition, in some embodiments, the sensor may be part of an
integrated sample acquisition and analyte measurement device. The
integrated sample acquisition and analyte measurement device may
include the sensor and a skin piercing member, so that the device
can be used to pierce the skin of a user to cause flow of a fluid
sample, such as blood, that can then be collected by the sensor. In
at least some embodiments, the fluid sample can be collected
without moving the integrated sample acquisition and analyte
measurement device.
One method of forming a sensor, as described above, includes
forming at least one working electrode on a first substrate and
forming at least one counter or counter/reference electrode on a
second substrate. A spacer layer is disposed on either the first or
second substrates. The spacer layer defines a channel into which a
sample can be drawn and held when the sensor is completed. A redox
mediator and/or second electron transfer agent are disposed on the
first or second substrate in a region that will be exposed within
the channel when the sensor is completed. The first and second
substrates are then brought together and spaced apart by the spacer
layer with the channel providing access to the at least one working
electrode and the at least one counter or counter/reference
electrode. In some embodiments, the first and second substrates are
portions of a single sheet or continuous web of material.
These and various other features which characterize the invention
are pointed out with particularity in the attached claims. For a
better understanding of the invention, its advantages, and
objectives obtained by its use, reference should be made to the
drawings and to the accompanying description, in which there is
illustrated and described preferred embodiments of the
invention.
BRIEF DESCRIPTION OF THE DRAWINGS
Referring now to the drawings, wherein like reference numerals and
letters indicate corresponding structure throughout the several
views:
FIG. 1 is a schematic view of a first embodiment of an
electrochemical sensor in accordance with the principles of the
present invention having a working electrode and a counter
electrode facing each other;
FIG. 2 is a schematic view of a second embodiment of an
electrochemical sensor in accordance with the principles of the
present invention having a working electrode and a counter
electrode in a coplanar configuration;
FIG. 3 is a schematic view of a third embodiment of an
electrochemical sensor in accordance with the principles of the
present invention having a working electrode and a counter
electrode facing each other and having an extended sample
chamber;
FIG. 4 is a not-to-scale side-sectional drawing of a portion of the
sensor of FIG. 1 or 3 showing the relative positions of the redox
mediator, the sample chamber, and the electrodes;
FIG. 5 is a top view of a fourth embodiment of an electrochemical
sensor in accordance with the principles of the present invention,
this sensor includes multiple working electrodes;
FIG. 6 is a perspective view of an embodiment of an analyte
measurement device, in accordance with the principles of the
present invention, having a sample acquisition means and the sensor
of FIG. 4;
FIG. 7 is a graph of the charge required to electrooxidize a known
quantity of glucose in an electrolyte buffered solution (filled
circles) or serum solution (open circles) using the sensor of FIG.
1 with glucose oxidase as the second electron transfer agent;
FIG. 8 is a graph of the average glucose concentrations for the
data of FIG. 7 (buffered solutions only) with calibration curves
calculated to fit the averages; a linear calibration curve was
calculated for the 10-20 mM concentrations and a second order
polynomial calibration curve was calculated for the 0-10 mM
concentrations;
FIG. 9 is a Clarke-type clinical grid analyzing the clinical
relevance of the glucose measurements of FIG. 7;
FIG. 10 is a graph of the charge required to electrooxidize a known
quantity of glucose in an electrolyte buffered solution using the
sensor of FIG. 1 with glucose dehydrogenase as the second electron
transfer agent;
FIGS. 11A, 11B, and 11C are top views of three configurations for
overlapping working and counter electrodes according to the present
invention;
FIGS. 12A and 1213 are cross-sectional views of one embodiment of
an electrode pair formed using a recess of a base material,
according to the invention;
FIGS. 13A and 13B are cross-sectional views of yet another
embodiment of an electrode pair of the present invention formed in
a recess of a base material;
FIGS. 14A and 14B are cross-sectional views of a further embodiment
of an electrode pair of the present invention formed using a recess
of a base material and a sorbent material;
FIG. 15 is a graph of charge delivered by a sensor having a
diffusible redox mediator over time for several concentrations of
glucose;
FIG. 16 is a graph of charge delivered by a sensor having a
diffusible redox mediator for several glucose concentrations;
FIG. 17 is a graph of charge delivered by sensors with different
amounts of diffusible redox mediator over time.
FIG. 18A illustrates a top view of a first film with a working
electrode for use in a fifth embodiment of a sensor according to
the invention;
FIG. 18B illustrates a top view of a spacer for placement on the
first film of FIG. 18A;
FIG. 18C illustrates a bottom view of a second film (inverted with
respect to FIGS. 18A and 18B) with counter electrodes placement
over the spacer of FIG. 18B and first film of FIG. 18A;
FIG. 19A illustrates a top view of a first film with a working
electrode for use in a sixth embodiment of a sensor according to
the invention;
FIG. 19B illustrates a top view of a spacer for placement on the
first film of FIG. 19A;
FIG. 19C illustrates a bottom view of a second film (inverted with
respect to FIGS. 19A and 1913) with counter electrodes placement
over the spacer of FIG. 19B and first film of FIG. 19A;
FIG. 20A illustrates a top view of a first film with a working
electrode for use in a seventh embodiment of a sensor according to
the invention;
FIG. 20B illustrates a top view of a spacer for placement on the
first film of FIG. 20A;
FIG. 20C illustrates a bottom view of a second film (inverted with
respect to FIGS. 20A and 20B) with counter electrodes placement
over the spacer of FIG. 20B and first film of FIG. 20A;
FIG. 21A illustrates a top view of a first film with a working
electrode for use in a eighth embodiment of a sensor according to
the invention;
FIG. 21B illustrates a top view of a spacer for placement on the
first film of FIG. 21A;
FIG. 21C illustrates a bottom view of a second film (inverted with
respect to FIGS. 21A and 21B) with counter electrodes placement
over the spacer of FIG. 21B and first film of FIG. 21A;
FIG. 22A illustrates a top view of a first film with a working
electrode for use in a ninth embodiment of a sensor according to
the invention;
FIG. 22B illustrates a top view of a spacer for placement on the
first film of FIG. 22A;
FIG. 22C illustrates a bottom view of a second film (inverted with
respect to FIGS. 22A and 22B) with counter electrodes placement
over the spacer of FIG. 22B and first film of FIG. 22A;
FIG. 23A illustrates a top view of a first film with a working
electrode for use in a tenth embodiment of a sensor according to
the invention;
FIG. 23B illustrates a top view of a spacer for placement on the
first film of FIG. 23A;
FIG. 23C illustrates a bottom view of a second film (inverted with
respect to FIGS. 23A and 23B) with counter electrodes placement
over the spacer of FIG. 23B and first film of FIG. 23A;
FIG. 24A illustrates a top view of a first film with a working
electrode for use in an eleventh embodiment of a sensor according
to the invention;
FIG. 24B illustrates a top view of a spacer for placement on the
first film of FIG. 24A;
FIG. 24C illustrates a bottom view of a second film (inverted with
respect to FIGS. 24A and 24B with counter electrodes placement over
the spacer of FIG. 24B and first film of FIG. 24A;
FIG. 25 illustrates a top view of a twelfth embodiment of an
electrochemical sensor, according to the invention;
FIG. 26 illustrates a perspective view of one embodiment of an
integrated analyte acquisition and sensor device;
FIG. 27 illustrates a cross-sectional view of a thirteenth
embodiment of a sensor, according to the invention;
FIG. 28 illustrates a graph comparing measurements of analyte
concentration in blood samples collected from a subject's arm made
by a sensor of the invention with those determined by a standard
blood test;
FIG. 29 illustrates a graph comparing measurements of analyte
concentration in blood samples collected from a subject's finger
made by a sensor of the invention with those determined by a
standard blood test;
FIG. 30 illustrates a graph comparing measurements of analyte
concentration in venous samples made by a sensor of the invention
with those determined by a standard blood test;
FIG. 31A illustrates a top view of one embodiment of a sheet of
sensor components, according to the invention;
FIG. 31B illustrates a top view of another embodiment of a sheet of
sensor components, according to the invention; and
FIG. 32 illustrates a cross-sectional view looking from inside the
meter to a sensor of the invention disposed in a meter.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT
When used herein, the following definitions define the stated
term:
An "air-oxidizable mediator" is a redox mediator that is oxidized
by air, preferably so that at least 90% of the mediator is in an
oxidized state upon storage in air either as a solid or as a liquid
during a period of time, for example, one month or less, and,
preferably, one week or less, and, more preferably, one day or
less.
"Amperometry" includes steady-state amperometry, chronoamperometry,
and Cottrell-type measurements.
A "biological fluid" is any body fluid in which the analyte can be
measured, for example, blood, interstitial fluid, dermal fluid,
sweat, and tears.
The term "blood" in the context of the invention includes whole
blood and its cell-free components, such as, plasma and serum.
"Coulometry" is the determination of charge passed or projected to
pass during complete or nearly complete electrolysis of the
analyte, either directly on the electrode or through one or more
electron transfer agents. The charge is determined by measurement
of charge passed during partial or nearly complete electrolysis of
the analyte or, more often, by multiple measurements during the
electrolysis of a decaying current and elapsed time. The decaying
current results from the decline in the concentration of the
electrolyzed species caused by the electrolysis.
A "counter electrode" refers to one or more electrodes paired with
the working electrode, through which passes an electrochemical
current equal in magnitude and opposite in sign to the current
passed through the working electrode. The term "counter electrode"
is meant to include counter electrodes which also function as
reference electrodes (i.e. a counter/reference electrode) unless
the description provides that a "counter electrode" excludes a
reference or counter/reference electrode.
An "effective diffusion coefficient" is the diffusion coefficient
characterizing transport of a substance, for example, an analyte,
an enzyme, or a redox mediator, in the volume between the
electrodes of the electrochemical cell. In at least some instances,
the cell volume may be occupied by more than one medium (e.g., the
sample fluid and a polymer film). Diffusion of a substance through
each medium may occur at a different rate. The effective diffusion
coefficient corresponds to a diffusion rate through this
multiple-media volume and is typically different than the diffusion
coefficient for the substance in a cell filled solely with sample
fluid.
An "electrochemical sensor" is a device configured to detect the
presence of and/or measure the concentration of an analyte via
electrochemical oxidation and reduction reactions. These reactions
are transduced to an electrical signal that can be correlated to an
amount or concentration of analyte.
"Electrolysis" is the electrooxidation or electroreduction of a
compound either directly at an electrode or via one or more
electron transfer agents (e.g., redox mediators and/or
enzymes).
The term "facing electrodes" refers to a configuration of the
working and counter electrodes in which the working surface of the
working electrode is disposed in approximate opposition to a
surface of the counter electrode. In at least some instances, the
distance between the working and counter electrodes is less than
the width of the working surface of the working electrode.
A compound is "immobilized" on a surface when it is entrapped on or
chemically bound to the surface.
An "indicator electrode" includes one or more electrodes that
detect partial or complete filling of a sample chamber and/or
measurement zone.
A "layer" includes one or more layers.
The "measurement zone" is defined herein as a region of the sample
chamber sized to contain only that portion of the sample that is to
be interrogated during an analyte assay.
A "non-diffusible," "non-leachable," or "non-releasable" compound
is a compound which does not substantially diffuse away from the
working surface of the working electrode for the duration of the
analyte assay.
The "potential of the counter/reference electrode" is the half cell
potential of the reference electrode or counter/reference electrode
of the cell when the solution in the cell is 0.1 M NaCl solution at
pH7.
"Potentiometry" and "chronopotentiometry" refer to taking a
potentiometric measurement at one or more points in time.
A "redox mediator" is an electron transfer agent for carrying
electrons between the analyte and the working electrode, either
directly, or via a second electron transfer agent.
A "reference electrode" includes a reference electrode that also
functions as a counter electrode (i.e., a counter/reference
electrode) unless the description provides that a "reference
electrode" excludes a counter/reference electrode.
A "second electron transfer agent" is a molecule that carries
electrons between a redox mediator and the analyte.
"Sorbent material" is material that wicks, retains, and/or is
wetted by a fluid sample and which typically does not substantially
prevent diffusion of the analyte to the electrode.
A "surface in the sample chamber" includes a surface of a working
electrode, counter electrode, counter/reference electrode,
reference electrode, indicator electrode, a spacer, or any other
surface bounding the sample chamber.
A "working electrode" is an electrode at which analyte is
electrooxidized or electroreduced with or without the agency of a
redox mediator.
A "working surface" is the portion of a working electrode that is
covered with non-leachable redox mediator and exposed to the
sample, or, if the redox mediator is diffusible, a "working
surface" is the portion of the working electrode that is exposed to
the sample.
The small volume, in vitro analyte sensors of the present invention
are designed to measure the concentration of an analyte in a
portion of a sample having a volume no more than about 1 .mu.L,
preferably no more than about 0.5 .mu.L, more preferably no more
than about 0.25 .mu.L, and most preferably no more than about 0.1
.mu.L. The analyte of interest is typically provided in a solution
or biological fluid, such as blood or serum. Referring to the
Drawings in general and FIGS. 1-4 in particular, a small volume, in
vitro electrochemical sensor 20 of the invention generally includes
a working electrode 22, a counter (or counter/reference) electrode
24, and a sample chamber 26 (see FIG. 4). The sample chamber 26 is
configured so that when a sample is provided in the chamber the
sample is in electrolytic contact with both the working electrode
22 and the counter electrode 24. This allows electrical current to
flow between the electrodes to effect the electrolysis
(electrooxidation or electroreduction) of the analyte.
Working Electrode
The working electrode 22 may be formed from a molded carbon fiber
composite or it may consist of an inert non-conducting base
material, such as polyester, upon which a suitable conducting layer
is deposited. The conducting layer typically has relatively low
electrical resistance and is typically electrochemically inert over
the potential range of the sensor during operation. Suitable
conducting layers include gold, carbon, platinum, ruthenium
dioxide, palladium, and conductive epoxies, such as, for example,
ECCOCOAT CT5079-3 Carbon-Filled Conductive Epoxy Coating (available
from W.R. Grace Company, Woburn, Mass.), as well as other
non-corroding materials known to those skilled in the art. The
electrode (e.g., the conducting layer) is deposited on the surface
of the inert material by methods such as vapor deposition or
printing.
A tab 23 may be provided on the end of the working electrode 22 for
easy connection of the electrode to external electronics (not
shown) such as a voltage source or current measuring equipment.
Other known methods or structures (such as contact pads) may be
used to connect the working electrode 22 to the external
electronics.
To prevent electrochemical reactions from occurring on portions of
the working electrode not coated by the mediator, when a
non-leachable mediator is used, a dielectric 40 may be deposited on
the electrode over, under, or surrounding the region with the redox
mediator, as shown in FIG. 4. Suitable dielectric materials include
waxes and non-conducting organic polymers such as polyethylene.
Dielectric 40 may also cover a portion of the redox mediator on the
electrode. The covered portion of the redox mediator will not
contact the sample, and, therefore, will not be a part of the
electrode's working surface.
Sensing Chemistry
In addition to the working electrode 22, sensing chemistry
materials are provided in the sample chamber 26 for the analysis of
the analyte. This sensing chemistry preferably includes a redox
mediator and a second electron transfer mediator, although in some
instances, one or the other may be used alone. The redox mediator
and second electron transfer agent can be independently diffusible
or non-leachable (i.e., non-diffusible) such that either or both
may be diffusible or non-leachable. Placement of sensor chemistry
components may depend on whether they are diffusible or
non-leachable. For example, non-leachable and/or diffusible
component(s) typically form a sensing layer on the working
electrode. Alternatively, one or more diffusible components may be
disposed on any surface in the sample chamber prior to the
introduction of the sample. As another example, one or more
diffusible component(s) may be placed in the sample prior to
introduction of the sample into the sensor.
If the redox mediator is non-leachable, then the non-leachable
redox mediator is typically disposed on the working electrode 22 as
a sensing layer 32. In an embodiment having a redox mediator and a
second electron transfer agent, if the redox mediator and second
electron transfer agent are both non-leachable, then both of the
non-leachable components are disposed on the working electrode 22
as a sensing layer 32.
If, for example, the second electron transfer agent is diffusible
and the redox mediator is non-leachable, then at least the redox
mediator is disposed on the working electrode 22 as a sensing layer
32. The diffusible second electron transfer agent need not be
disposed on a sensing layer of the working electrode, but may be
disposed on any surface of the sample chamber, including within the
redox mediator sensing layer, or may be placed in the sample. If
the redox mediator is diffusible, then the redox mediator may be
disposed on any surface of the sample chamber or may be placed in
the sample.
If both the redox mediator and second electron transfer agent are
diffusible, then the diffusible components may be independently or
jointly disposed on any surface of the sample chamber and/or placed
in the sample (i.e., each diffusible component need not be disposed
on the same surface of the sample chamber or placed in the
sample).
The redox mediator, whether it is diffusible or non-leachable,
mediates a current between the working electrode 22 and the analyte
and enables the electrochemical analysis of molecules which may not
be suited for direct electrochemical reaction on an electrode. The
mediator functions as an electron transfer agent between the
electrode and the analyte.
In one embodiment, the redox mediator and second electron transfer
agent are diffusible and disposed on the same surface of the sample
chamber, such as, for example, on the working electrode. In this
same vein, both can be disposed on, for example, the counter
electrode, counter/reference electrode, reference electrode, or
indicator electrode. In other instances, the redox mediator and
second electron transfer agent are both diffusible and
independently placed on a surface of the sample chamber and/or in
the sample. For example, the redox mediator may be placed on the
working electrode while the second electron transfer agent is
placed on any surface, except for the working electrode, or is
placed in the sample. Similarly, the reverse situation in which the
second electron transfer agent is disposed on the working electrode
and the redox mediator is disposed on any surface, except for the
working electrode, or is placed in the sample is also a suitable
embodiment. As another example, the redox mediator may be disposed
on the counter electrode and the second electron transfer agent is
placed on any surface except for the counter electrode or is placed
in the sample. The reverse situation is also suitable.
The diffusible redox mediator and/or second electron transfer agent
may diffuse rapidly into the sample or diffusion may occur over a
period of time. Similarly, the diffusible redox mediator and/or
second electron transfer agent may first dissolve from the surface
on which it was placed as a solid and then the diffusible redox
mediator and/or second electron transfer agent may diffuse into the
sample, either rapidly or over a period of time. If the redox
mediator and/or second electron transfer agent diffuse over a
period of time, a user may be directed to wait a period of time
before measuring the analyte concentration to allow for diffusion
of the redox mediator and/or second electron transfer agent.
Background Signal
In at least some instances, a diffusible redox mediator may shuttle
back and forth from the working electrode to the counter electrode
even in the absence of analyte. This typically creates a background
signal. For coulometric measurements, this background signal is
referred to herein as "Q.sub.Back." The background signal
corresponds to the charge passed in an electrochemical assay in the
absence of the analyte. The background signal typically has both a
transient component and a steady-state component. At least a
portion of the transient component may result, for example, from
the establishment of a concentration gradient of the mediator in a
particular oxidation state. At least a portion of the steady-state
component may result, for example, from the redox mediator
shuttling between the working electrode and counter or
counter/reference electrode. Shuttling refers to the redox mediator
being electrooxidized (or electroreduced) at the working electrode
and then being electroreduced (or electrooxidized) at the counter
or counter/reference electrode, thereby making the redox mediator
available to be electrooxidized (or electroreduced) again at the
working electrode so that the redox mediator is cycling between
electrooxidation and electroreduction.
The amount of shuttling of the redox mediator, and therefore, the
steady-state component of the background signal varies with, for
example, the effective diffusion coefficient of the redox mediator,
the viscosity of the sample, the temperature of the sample, the
dimensions of the electrochemical cell, the distance between the
working electrode and the counter or counter/reference electrode,
and the angle between the working electrode and the counter or
counter/reference electrode.
In some instances, the steady-state component of the background
signal may contain noise associated with (a) variability in, for
example, the temperature of the sample, the sample viscosity, or
any other parameter on which the background signal depends during
the duration of the assay, or (b) imperfections in the
electrochemical cell, such as, for example, non-uniform separation
between the working electrode and the counter or counter/reference
electrode, variations in electrode geometry, or protrusions from
the working electrode, the counter electrode, and/or the
counter/reference electrode.
Although the steady-state component of the background signal may be
reproducible, any noise inherently is not reproducible. As a
result, the noise adversely affects accuracy. In some cases, the
background signal and noise are related. As a result, the noise,
and the error it introduces, can be reduced by reducing the
background signal. For example, reducing the shuttling of the
mediator between the working electrode and counter electrode or
counter/reference electrode will likely reduce the noise associated
with changes in sample temperature and viscosity which affect
diffusion of the redox mediator.
Thus, to increase the accuracy of the measurements or to decrease
error in the measurements in those instances when reducing a
background signal also reduces noise, a moderate to near-zero level
of background signal is desirable. In at least some instances, the
sensor is constructed so that the background signal is at most five
times the size of a signal generated by electrolysis of an amount
of analyte. Preferably, the background signal is at most 200%,
100%, 50%, 25%, 10%, or 5% of the signal generated by electrolysis
of the analyte. In the case of amperometry, this comparison may be
made by determining the ratio of the current from the shuttling of
the redox mediator to the current generated by the electrolysis of
the analyte. In the case of potentiometry, this comparison may be
made by determining the potential measurement from the shuttling of
the redox mediator and the potential measurement generated by
electrolysis of the analyte. In the case of coulometry, this
comparison may be made by determining the charge transferred at the
working electrode by the shuttling of the redox mediator and the
charge transferred at the working electrode by the electrolysis of
the analyte.
The size of the background signal may be compared to a
predetermined amount of analyte. The predetermined amount of
analyte in a sample may be, for example, an expected or average
molar amount of analyte. The expected or average molar amount of
analyte may be determined as, for example, the average value for
users or individuals; an average value for a population; a maximum,
minimum, or average of a normal physiological range; a maximum or
minimum physiological value for a population; a maximum or minimum
physiological value for users or individuals; an average, maximum,
or minimum deviation outside a normal physiological range value for
users, individuals, or a population; a deviation above or below an
average value for a population; or an average, maximum, or minimum
deviation above or below an average normal physiological value for
users or individuals. A population may be defined by, for example,
health, sex, or age, such as, for example, a normal adult, child,
or newborn population. If a population is defined by health, the
population may include people lacking a particular condition or
alternatively, having a particular condition, such as, for example,
diabetes. Reference intervals pertaining to average or expected
values, such as, for example, those provided in Tietz Textbook of
Clinical Chemistry, Appendix (pp. 2175-2217) (2nd Ed., Carl A.
Burtis and Edward R. Ashwood, eds., W.D. Saunders Co., Philadelphia
1994) (incorporated herein by reference) may be used as guidelines,
but a physical examination or blood chemistry determination by a
skilled physician may also be used to determine an average or
expected value for an individual. For example, an adult may have
glucose in a concentration of 65 to 95 mg/dL in whole blood or
L-lactate in a concentration of 8.1 to 15.3 mg/dL in venous whole
blood after fasting, according to Tietz Textbook of Clinical
Chemistry. An average normal physiological concentration for an
adult, for example, may then correspond to 80 mg/dL for glucose or
12.7 mg/dL for lactate. Other examples include a person having
juvenile onset diabetes, yet good glycemic control, and a glucose
concentration between about 50 mg/dL and 400 mg/dL thereby having
an average molar amount of 225 mg/dL. In another instance, a
non-diabetic adult may have a glucose concentration between about
80 mg/dL (after fasting) and 140 mg/dL (after consuming food),
thereby having an average molar amount of 110 mg/dL.
Additional analytes that may be determined include, for example,
acetyl choline, amylase, bilirubin, cholesterol, chorionic
gonadotropin, creatine kinase (e.g., CK-MB), creatine, DNA,
fructosamine, glucose, glutamine, growth hormones, hormones,
ketones, lactate, peroxide, prostate-specific antigen, prothrombin,
RNA, thyroid stimulating hormone, and troponin. The concentration
of drugs, such as, for example, antibiotics (e.g., gentamicin,
vancomycin, and the like), digitoxin, digoxin, drugs of abuse,
theophylline, and warfarin, may also be determined. Assays suitable
for determining the concentration of DNA and/or RNA are disclosed
in U.S. Pat. No. 6,281,006, U.S. patent application having Ser. No.
09/145,776 and described in U.S. Provisional Patent Application
Ser. Nos. 60/090,517, 60/093,100, and 60/114,919, incorporated
herein by reference.
To construct a sensor having a particular ratio of background
signal to analyte signal from electrolysis, several parameters
relating to current and/or charge from the redox mediator shuttling
background signal and/or from the signal generated by electrolysis
of the analyte may be considered and chosen to obtain a desired
ratio. Typically, the signal determined for a coulometric assay is
the charge; whereas the signal determined for an amperometric assay
is the current at the time when the measurement is taken. Because
the current and charge depend on several parameters, the desired
ratio for background signal generated by shuttling of the redox
mediator to signal generated by electrolysis of the analyte may be
accomplished by a variety of sensor configurations and methods for
operating a sensor.
Controlling Background Signal
One method of controlling background signal includes using a redox
mediator that a) oxidizes the analyte at a half wave potential, as
measured by cyclic voltammetry in 0.1 M NaCl at pH 7, of no more
than about +100 mV relative to the potential of a reference or
counter/reference electrode or b) reduces the analyte at a half
wave potential, as measured by cyclic voltammetry in 0.1 M NaCl at
pH 7, of no less than about -100 mV relative to the potential of a
reference or counter/reference electrode. A suitable reference or
counter/reference electrode (e.g., a silver/silver chloride
electrode) may be chosen. Preferably, the redox mediator a)
oxidizes the analyte at a half wave potential, as measured by
cyclic voltammetry in 0.1 M NaCl at pH 7, of no more than about +50
mV, +25 mV, 0 mV, -25 mV, -50 mV, -100 mV, or -150 mV relative to
the potential of the reference or counter/reference electrode or b)
reduces the analyte at a half wave potential, as measured by cyclic
voltammetry in 0.1 M NaCl at pH 7, of no less than about -50 mV,
-25 mV, 0 mV, +25 mV, +50 mV, +100 mV, +150 mV, or +200 mV relative
to the potential of the reference or counter/reference electrode.
Alternatively, in the case of reduction of the redox mediator by
the counter electrode, the sensor is operated at an applied
potential of no more than about +100 mV, +50 mV, +25 mV, 0 mV, -25
mV, -50 mV, -100 mV, or -150 mV between the working electrode and
the counter or counter/reference electrode. In the case of
oxidation of the redox mediator at the counter electrode, the
sensor is operated at an applied potential of no less than about
-100 mV, -50 mV, -25 mV, 0 mV, +25 mV, +50 mV, +100 mV, +150 mV, or
+200 mV between the working electrode and the counter or
counter/reference electrode.
Another method includes controlling the applied potential such that
for an electrooxidative assay the redox mediator is not readily
reduced at the counter or counter/reference electrode or for an
electroreductive assay the redox mediator is not readily oxidized
at the counter or counter/reference electrode. This can be
accomplished, for example, in an electrooxidative assay by using a
sensor having a diffusible redox mediator with a potential,
relative to a reference or counter/reference electrode, that is
negative with respect to the potential of the counter electrode
(relative to a reference electrode) or the counter/reference
electrode. The potential (relative to a reference or
counter/reference electrode) of the working electrode is chosen to
be positive with respect to the redox mediator and may be negative
with respect to the counter or counter/reference electrode, so that
the redox mediator is oxidized at the working electrode. For
example, when the electrooxidation of an analyte is mediated by a
diffusible redox mediator with a potential of -200 mV versus the
reference or counter/reference electrode, and the potential at
which the working electrode is poised is -150 mV relative to the
reference or counter/reference electrode, then the redox mediator
is substantially oxidized at the working electrode and will oxidize
the analyte. Further, if some of the oxidized redox mediator
reaches the counter or counter/reference electrode, the redox
mediator will not be readily reduced at the counter or
counter/reference electrode because the counter or
counter/reference electrode is poised well positive (i.e., 150 mV)
of the potential of the redox mediator.
In an electroreductive assay, a sensor is provided having a
diffusible redox mediator with a formal potential, relative to a
reference or counter/reference electrode, that is positive with
respect to the potential of the counter or counter/reference
electrode. The potential, relative to the reference or
counter/reference electrode, of the working electrode is chosen to
be negative with respect to the redox mediator and may be poised
positive with respect to the counter or counter/reference
electrode, so that the redox mediator is reduced at the working
electrode.
Still another method of limiting background current includes having
the redox mediator become immobilized when reacted on the counter
electrode or counter/reference electrode by, for example,
precipitation or polymerization. For example, the mediator may be
cationic in the oxidized state, but neutral and much less soluble
in the reduced state. Reduction at the counter/reference electrode
leads to precipitation of the reduced, neutral mediator on the
counter/reference electrode.
Another sensor configuration suitable for controlling background
signal includes a sensor having a molar amount of redox mediator
that is stoichiometrically the same as or less than an expected or
average molar amount of analyte. The expected or average molar
amount of analyte may be determined as already explained above. The
expected or average molar amount of analyte may be determined as,
for example, the average value for users or individuals; an average
value for a population; a maximum, minimum, or average of a normal
physiological range; a maximum or minimum physiological value for a
population; a maximum or minimum physiological value for users or
individuals; an average, maximum, or minimum deviation outside a
normal physiological range value for users, individuals, or a
population; a deviation above or below an average value for a
population; or an average, maximum, or minimum deviation above or
below an average normal physiological value for users or
individuals. A population may be defined by, for example, health,
sex, or age, such as, for example, a normal adult, child, or
newborn population. If a population is defined by health, the
population may include people lacking a particular condition or
alternatively, having a particular condition, such as, for example,
diabetes. Reference intervals pertaining to average or expected
values, such as, for example, those provided in Tietz Textbook of
Clinical Chemistry, supra, may be used as guidelines, but a
physical examination or blood chemistry determination may also
determine an average or expected value. For example, the
physiological average molar amount of analyte may depend on the
health or age of the person from whom the sample is obtained. This
determination is within the knowledge of a skilled physician.
By reducing the concentration of the redox mediator relative to the
concentration of the analyte, the signal attributable to the
analyte relative to the signal attributable to the shuttling of the
redox mediator is increased. In implementation of this method, the
molar amount of redox mediator may be no more than 50%, 20%, 10%,
or 5%, on a stoichiometric basis, of the expected or average molar
amount of analyte.
The amount of redox mediator used in such a sensor configuration
should fall within a range. The upper limit of the range may be
determined based on, for example, the acceptable maximum signal due
to shuttling of the redox mediator; the design of the
electrochemical cell, including, for example, the dimensions of the
cell and the position of the electrodes; the effective diffusion
coefficient of the redox mediator; and the length of time needed
for the assay. Moreover, the acceptable maximum signal due to redox
mediator shuttling may vary from assay to assay as a result of one
or more assay parameters, such as, for example, whether the assay
is intended to be qualitative, semi-quantitative, or quantitative;
whether small differences in analyte concentration serve as a basis
to modify therapy; and the expected concentration of the
analyte.
Although it is advantageous to minimize the amount of redox
mediator used, the range for the acceptable amount of redox
mediator does typically have a lower limit. The minimum amount of
redox mediator that may be used is the concentration of redox
mediator that is necessary to accomplish the assay within a
desirable measurement time period, for example, no more than about
5 minutes or no more than about 1 minute. The time required to
accomplish an assay depends on, for example, the distance between
the working electrode and the counter or counter/reference
electrode, the effective diffusion coefficient of the redox
mediator, and the concentration of the analyte. In some instances,
for example, when no kinetic limitations are present, i.e.,
shuttling of the redox mediator depends only on diffusion, the
minimum concentration of redox mediator may be determined by the
following formula: C.sub.m=(d.sup.2C.sub.A)/D.sub.mt where C.sub.m
is the minimum concentration of mediator required; d is the
distance between a working electrode and a counter or
counter/reference electrode in a facing arrangement; C.sub.A is the
average analyte concentration in the sample; D.sub.m is the
effective diffusion coefficient of the mediator in the sample; and
t is the desired measurement time.
For example, when the distance between the facing electrode pair is
50 the analyte being measured is 5 mM glucose, the redox mediator
effective diffusion coefficient is 10.sup.-6 cm.sup.2/sec and the
desirable response time is no more than about 1 minute, then the
minimum redox mediator concentration is 2.08 mM. Under these
conditions the background signal will be less than the signal from
the electrooxidation of the analyte.
Yet another sensor configuration for limiting the background
current generated by a diffusible redox mediator includes having a
barrier to the flow of the diffusible mediator to the counter
electrode. The barrier can be, for example, a film through which
the redox mediator can not diffuse or through which the redox
mediator diffuses slowly. Examples of suitable films include
polycarbonate, polyvinyl alcohol, and regenerated cellulose or
cellulose ester membranes. Alternatively, the barrier can include
charged or polar particles, compounds, or functional groups to
prevent or reduce the flow of a charged redox mediator relative to
the flow of a charge neutral or less charged analyte. If the redox
mediator is positively charged, as are many of the osmium redox
mediators described below, the barrier can be a positively charged
or polar film, such as a methylated poly(1-vinyl imidazole). If the
redox mediator is negatively charged, the barrier can be a
negatively charged or polar film, such as Nafion.RTM.. Examples of
suitable polar matrices include a bipolar membrane, a membrane
having a cationic polymer cross-linked with an anionic polymer, and
the like. In some instances, the barrier reduces the oxidation or
reduction of the diffusible redox mediator at the counter electrode
by at least 25%, 50%, or 90%.
Still another sensor configuration for limiting the background
current includes a sensor having a redox mediator that is more
readily oxidized or reduced on the working electrode than reduced
or oxidized on the counter electrode. The rate of reaction of the
redox mediator at an electrode can be a function of the material of
the electrode. For example, some redox mediators may react faster
at a carbon electrode than at a Ag/AgCl electrode. Appropriate
selection of the electrodes may provide a reaction rate at one
electrode that is significantly slower than the rate at the other
electrode. In some instances, the rate of oxidation or reduction of
the diffusible redox mediator at the counter electrode is reduced
by at least 25%, 50%, or 90%, as compared to the working electrode.
In some instances the rate of reaction for the redox mediator at
the counter or counter/reference electrode is controlled by, for
example, choosing a material for the counter or counter/reference
electrode that would require an overpotential or a potential higher
than the applied potential to increase the reaction rate at the
counter or counter/reference electrode.
Another sensor configuration for limiting background current
includes elements suitable for reducing the diffusion of the redox
mediator. Diffusion can be reduced by, for example, using a redox
mediator with a relatively low diffusion coefficient or increasing
the viscosity of the sample in the measurement zone. In another
embodiment, the diffusion of the redox mediator may be decreased by
choosing a redox mediator with high molecular weight, such as, for
example, greater than 5,000 daltons, preferably greater than 25,000
daltons, and more preferably greater than 100,000 daltons.
Redox Mediators
Although any organic or organometallic redox species can be used as
a redox mediator, one type of suitable redox mediator is a
transition metal compound or complex. Examples of suitable
transition metal compounds or complexes include osmium, ruthenium,
iron, and cobalt compounds or complexes. In these complexes, the
transition metal is coordinatively bound to one or more ligands.
The ligands are typically mono-, di-, tri-, or tetradentate. The
most preferred ligands are heterocyclic nitrogen compounds, such
as, for example, pyridine and/or imidazole derivatives.
Multidentate ligands may include multiple pyridine and/or imidazole
rings. Alternatively, metallocene derivatives, such as, for
example, ferrocene, can be used.
Suitable redox mediators include osmium or ruthenium transition
metal complexes with one or more ligands, each ligand having one or
more nitrogen-containing heterocycles. Examples of such ligands
include pyridine and imidazole rings and ligands having two or more
pyridine and/or imidazole rings such as, for example,
2,2'-bipyridine; 2,2':6',2''-terpyridine; 1,10-phenanthroline; and
ligands having the following structures:
##STR00001## and derivatives thereof, wherein R.sub.1 and R.sub.2
are each independently hydrogen, hydroxy, alkoxy, alkenyl, vinyl,
allyl, amido, amino, vinylketone, keto, or sulfur-containing
groups.
The term "alkyl" includes a straight or branched saturated
aliphatic hydrocarbon chain having from 1 to 6 carbon atoms, such
as, for example, methyl, ethyl isopropyl (1-methylethyl), butyl,
tert-butyl (1,1-dimethylethyl), and the like. Preferably the
hydrocarbon chain has from 1 to 3 carbon atoms.
The term "alkoxy" includes an alkyl as defined above joined to the
remainder of the structure by an oxygen atom, such as, for example,
methoxy, ethoxy, propoxy, isopropoxy (1-methylethoxy), butoxy,
tert-butoxy, and the like.
The term "alkenyl" includes an unsaturated aliphatic hydrocarbon
chain having from 2 to 6 carbon atoms, such as, for example,
ethenyl, 1-propenyl, 2-propenyl, 1-butenyl, 2-methyl-1-propenyl,
and the like. Preferably the hydrocarbon chain has from 2 to 3
carbon atoms.
The term "amido" includes groups having a nitrogen atom bonded to
the carbon atom of a carbonyl group and includes groups having the
following formulas:
##STR00002## wherein R.sub.3 and R.sub.4 each independently
hydrogen, alkyl, alkoxy, or alkenyl.
The term "amino" as used herein includes alkylamino, such as
methylamino, diethylamino, N,N-methylethylamino and the like;
alkoxyalkylamino, such as N-(ethoxyethyl)amino,
N,N-di(methoxyethyl)amino, N,N-(methoxyethyl)(ethoxyethly)amino,
and the like; and nitrogen containing rings, such as piperidino,
piperazino, morpholino, and the like.
The term "vinylketone" includes a group having the formula:
##STR00003## wherein R.sub.5, R.sub.6, and R.sub.7 are each
independently hydrogen, alkyl, alkoxy, or alkenyl.
The term "keto" includes a group having the formula:
##STR00004## wherein R.sub.8 is hydrogen, alkyl, alkoxy, or
alkenyl.
The term "sulfur-containing group" includes mercapto, alkylmercapto
(such as methylmercapto, ethylmercapto, and the like),
alkoxyalkylmercapto (such as methoxyethylmercapto and the like),
alkoxyalkylsulfoxide (such as ethoxyethylsulfoxide and
propylsulfoxide and the like), alkoxyalkylsulfoxide (such as
ethoxyethylsulfoxide and the like), alkylsulfone (such as
methoxyethylsulfone and the like), and alkoxyalkylsulfone (such as
methoxyethylsulfone and the like). Preferably, the
sulfur-containing group is a mercapto group.
Other suitable redox mediators include osmium or ruthenium
transition metal complexes with one or more ligands, each ligand
having one or more nitrogen-containing heterocycles and each
nitrogen-containing heterocycle having a second heteroatom selected
from the group consisting of nitrogen, oxygen, sulfur, and
selenium.
Examples of ligands having one or more nitrogen-containing
heterocycles and in which each heterocycle has a second heteroatom
include ligands having the following structures:
##STR00005## wherein Y.sub.1, Y.sub.2, Y.sub.3, and Y.sub.4 are
each independently an oxygen atom, a sulfur atom, a selenium atom,
or a substituted nitrogen atom having the formula NR.sub.9 wherein
R.sub.9 is hydrogen, hydroxy, alkyl, alkoxy, alkenyl, amido, amino,
vinylketone, keto, or sulfur-containing group. The terms "alkyl,"
"alkoxy" "alkenyl," "amido," "amino," "vinylketone," "keto," and
"sulfur-containing group" are as defined above.
Suitable derivatives of these ligands include, for example, the
addition of alkyl, alkoxy, alkenyl, vinylester, and amido
functional groups to any of the available sites on the heterocyclic
ring, including, for example, on the 4-position (i.e., para to
nitrogen) of the pyridine rings or on one of the nitrogen atoms of
the imidazole ring.
Suitable derivatives of these ligands include, for example, the
addition of alkyl, alkoxy, alkenyl, vinylester, and amido
functional groups to any of the available sites on the heterocyclic
ring, including, for example, on the 4-position (i.e., para to
nitrogen) of the pyridine rings or on one of the nitrogen atoms of
the imidazole ring.
Suitable derivatives of 2,2'-bipyridine for complexation with the
osmium cation include, for example, mono-, di-, and
polyalkyl-2,2'-bipyridines, such as 4,4'-dimethyl-2,2'-bipyridine;
mono-, di-, and polyalkoxy-2,2'-bipyridines, such as
4,4'-dimethoxy-2,2'-bipyridine and 2,6'-dimethoxy-2,2'-bipyridine;
mono-, di-, and polyacetamido-2,2'-bipyridines, such as
4,4'-di(acetamido)-2,2'-bipyridine; mono-, di-, and
polyalkylaminoalkoxy-2,2'-bipyridines, such as
4,4'-di(N,N-dimethylaminoethoxy)-2,2'-bipyridine; and substituted
mono-, di-, and polypyrazolyl-2,2'-bipyridines, such as
4,4'-dimethoxy-6-(N-pyrazolyl)-2,2'-bipyridine and
4,4'-dimethoxy-6-(N-pyrazolylmethyl)-2,2'-bipyridine.
Suitable derivatives of 1,10-phenanthroline for complexation with
the osmium cation include, for example, mono-, di-, and
polyalkyl-1,10-phenanthrolines, such as
4,7-dimethyl-1,10-phenanthroline and mono, di-, and
polyalkoxy-1,10-phenanthrolines, such as
4,7-dimethoxy-1,110-phenanthroline and
5-methoxy-1,10-phenanthrolines.
Suitable derivatives for 2,2':6',2''-terpyridine include, for
example, mono-, di-, tri-, and polyalkyl-2,2':6',2''-terpyridines,
such as 4,4',4''-trimethyl-2,2':6',2''terpyridine,
4,4',4''-triethyl-2,2':6',2''-terpyridine, and mono-, di-, tri-,
and polyalkoxy-2,2':6',2''-terpyridines, such as
4,4',4''-trimethoxy-2,2':6',2''-terpyridine and
4'-methoxy-2,2':6',2''-terpyridine, and mono-, di-, tri-, and
polyamino-2,2':6',2''-terpyridine, such as
4'-amino-2,2':6',2''-terpyridine, and mono-, di-, tri-, and
polyalkylamino-2,2':6',2''-terpyridine, such as
4'-dimethylamino-2,2':6',2''-terpyridine, and mono-, di-, tri-, and
polyalkylthio-2,2':6',2''-terpyridine such as
4'-methylthio-2,2':6',2'-terpyridine and
4-methylthio-4'-ethylthio-2,2':6,2''-terpyridine.
Suitable derivatives for pyridine include, for example, mono-, di-,
tri-, and polysubstituted pyridines, such as
2,6-bis(N-pyrazolyl)pyridine,
2,6-bis(3-methyl-N-pyrazolyl)pyridine,
2,6-bis(2-imidazolyl)pyridine,
2,6-bis(1-methyl-2-imidazolyl)pyridine, and
2,6-bis(1-vinyl-2-imidazolyl)pyridine, and mono-, di-, tri-, and
polyaminopyridines, such as 4-aminopyridine,
4,4'-diaminobipyridine, 4,4'-di(dimethylamino)bipyridine, and
4,4',4''-triamino terpyridine.
Other suitable derivatives include compounds comprising three
heterocyclic rings. For example, one suitable derivative includes a
compound of the formula:
##STR00006## wherein R.sub.10, R.sub.11, and R.sub.12 are each
independently hydrogen, hydroxy, alkyl, alkoxy, alkenyl, vinyl,
allyl, amido, amino, vinylketone, keto, or sulfur-containing
group.
The terms "alkyl," "alkoxy." "alkenyl," "amido," "amino,"
"vinylketone," "keto," and "sulfur-containing group" are as defined
above.
Other suitable redox mediator derivatives include compounds of the
formula:
##STR00007## wherein R.sub.3 is hydrogen, hydroxy, alkyl, alkoxy,
alkenyl, vinyl, allyl, vinylketone, keto, amido, amino, or
sulfur-containing group; and Y.sub.5 and Y.sub.6 are each
independently a nitrogen or carbon atom.
The terms "alkyl," "alkoxy," "alkenyl," "amido," "amino,"
"vinylketone," "keto," and "sulfur-containing group" are as defined
above.
Still other suitable derivatives include compounds of the
formula:
##STR00008## wherein R.sub.14 is as defined above and Y.sub.7 and
Y.sub.8 are each independently a sulfur or oxygen atom.
Examples of suitable redox mediators also include, for example,
osmium cations complexed with (a) two bidentate ligands, such as
2,2'-bipyridine, 1,10-phenanthroline, or derivatives thereof (the
two ligands not necessarily being the same), (b) one tridentate
ligand, such as 2,2',2''-terpyridine and
2,6-di(imidazol-2-yl)pyridine, or (c) one bidentate ligand and one
tridentate ligand. Suitable osmium transition metal complexes
include the example, [(bpy).sub.2OsLX].sup.+/2+,
[(dimet).sub.2OsLX].sup.+/2+, [(dmo).sub.2OsLX].sup.+/2+,
[terOsLX.sub.2].sup.0/+, [trimetOsLX.sub.2].sup.0/+, and
[(ter)(bpy)LOs].sup.2+/3- where bpy is 2,2'-bipyridine, dimet is
4,4'-dimethyl-2,2'-bipyridine, dmo is
4,4'-dimethoxy-2,2'-bipyridine, ter is 2,2':6',2''-terpyridine,
trimet is 4,4',4''-trimethyl-2,2':6',2''-terpyridine, L is a
nitrogen-containing heterocyclic ligand, and X is a halogen, such
as fluorine, chlorine, or bromine.
The redox mediators often exchange electrons rapidly with each
other and with the electrode so that the complex can be rapidly
oxidized and/or reduced. In general, iron complexes are more
oxidizing than ruthenium complexes, which, in turn, are more
oxidizing than osmium complexes. In addition, the redox potential
generally increases with the number of coordinating heterocyclic
rings; six-membered heterocyclic rings increase the potential more
than five membered rings, except when the nitrogen coordinating the
metal is formally an anion. This is the case only if the nitrogen
in the ring is bound to both of its neighboring carbon atoms by
single bonds. If the nitrogen is formally an anion then the redox
potential generally increases more upon coordination of the metal
ion.
At least some diffusible redox mediators include one or more
pyridine or imidazole functional groups. The imidazole functional
group can also include other substituents and can be, for example,
vinyl imidazole, e.g., 1-vinyl imidazole, or methylimidazole, e.g.,
1-methylimidazole. Examples of suitable diffusible mediators may
include [Os(dmo).sub.2 (1-vinyl imidazole)X]X, [Os(dmo).sub.2
(1-vinyl imidazole)X]X.sub.2, [Os(dmo).sub.2 (imidazole)X]X,
[Os(dmo).sub.2 (imidazole)X]X.sub.2, [Os(dmo).sub.2
(1-methylimidazole)X]X.sub.2, and [Os(dmo).sub.2
(methylimidazole)X]X.sub.2, where dmo is
4,4'-dimethoxy-2,2'-bipyridine, and X is halogen as described
above.
Other osmium-containing redox mediators include
[Os((methoxy).sub.2phenanthroline).sub.2(N-methylimidazole)X].sup.+/2+;
[Os((acetamido).sub.2bipyridine).sub.2(L)X].sup.+/2+, where L is a
monodentate nitrogen-containing compound (including, but not
limited to, an imidazole derivative) chosen to refine the
potential; and Os(terpyridine)(L).sub.2Cl, where L is an
aminopyridine, such as a dialkylaminopyridine; an N-substituted
imidazole, such as N-methyl imidazole; an oxazole; a thiazole; or
an alkoxypyridine, such as methoxypyridine. X is halogen as
described above.
Osmium-free diffusible redox mediators include, for example,
phenoxazines, such as, 7-dimethylamino-1,2-benzophenoxazine
(Meldola Blue), 1,2-benzophenoxazine, and Nile Blue;
3-.beta.-naphthoyl (Brilliant Cresyl Blue);
tetramethylphenylcnediamine (TMPD); dichlorophenolindophenol
(DCIP); N-methyl phenazonium salts, for example, phenazine
methosulfate (PMS), N-methylphenazine methosulfate and
methoxyphenazine methosulfate; tetrazolium salts, for example,
tetrazolium blue or nitrotetrazolium blue; and phenothiazines, for
example, toluidine blue O.
Examples of other redox species include stable quinones and species
that in their oxidized state have quinoid structures, such as Nile
Blue and indophenol. Examples of suitable quinones include, for
example, derivatives of naphthoquinone, phenoquinone, benzoquinone,
naphthenequinone, and the like. Examples of naphthoquinone
derivatives include juglone (i.e., 5-hydroxy-1,4-naphthoquinone)
and derivatives thereof, such as, for example,
2,3-dichloro-5,8-dihydroxy-1,4-naphthoquinone,
2,3-dimethyl-5,8-dihydroxy-1,4-naphthoquinone,
2-chloro-5,8-dihydroxy-1,4-naphthoquinone,
2,3-methoxy-5-hydroxy-1,4-naphthoquinone, and the like. Other
examples include aminonaphthoquinones, such as, for example,
morpholino-naphthoquinones, such as
2-chloro-3-morpholino-1,4-naphthoquinone;
piperidino-naphthoquinones, such as
2-methyl-3-peperidino-1,4-naphthoquinone;
piperazino-naphthoquinones, such as
2-ethoxy-3-piperazino-1,4-naphthoquinone; and the like.
Suitable phenoquinone derivatives include, for example,
coerulignone (i.e., 3,3',5,5'-tetramethoxydiphenoquinone) and
derivatives thereof, such as, for example,
3,3',5,5'-tetramethyldiphenoquinone,
3,3',5,5'-tetrahydroxydiphenoquinone, and the like.
Suitable benzoquinone derivatives include, for example, coenzyme
Q.sub.0 (i.e., 2,3-dimethoxy-5-methyl-1,4-benzoquinone) and
derivatives thereof, such as, for example,
2,3,5-trimethyl-1,4-benzoquinone,
2,3-dimethyl-5-methoxy-1,4-benzoquinone,
2,3-dimethyl-5-hydroxy-1,4-benzoquinone, and the like.
Other suitable quinone derivatives include, for example,
acenaphthenequinone and ubiquinones, such as, for example, coenzyme
Q, including Q.sub.1, Q.sub.2, Q.sub.6, Q.sub.7, Q.sub.9, and
Q.sub.10.
Still other suitable osmium-free diffusible redox mediators
include, for example, Taylor's blue (i.e., 1,9-dimethylmethylene
blue), N,N'-diethylthiacyanine iodide, and thionine.
In another method, a sensing layer 32 contains a non-leachable
(i.e., non-releasable) redox mediator and is disposed on a portion
of the working electrode 22. The non-leachable redox mediator can
be, for example, a redox polymer (i.e., a polymer having one or
more redox species). Preferably, there is little or no leaching of
the non-leachable redox mediator away from the working electrode 22
into the sample during the measurement period, which is typically
less than about 5 minutes. The redox mediators of this embodiment
can be bound or otherwise immobilized on the working electrode 22
to prevent leaching of the mediator into the sample. The redox
mediator can be bound or otherwise immobilized on the working
electrode by known methods, for example, formation of multiple ion
bridges with a countercharged polyelectrolyte, covalent attachment
of the redox mediator to a polymer on the working electrode,
entrapment of the redox mediator in a matrix that has a high
affinity for the redox mediator, or bioconjugation of the redox
mediator with a compound bound to the working electrode. In one
embodiment, a cationic exchange membrane may be used to entrap an
anionic redox compound. Similarly, in another embodiment, an
anionic exchange membrane may be used to entrap a cationic redox
compound. In still another embodiment involving bioconjugation, a
biotin-bound redox mediator can conjugate with avidin or
straptavidin in a matrix near or immobilized on the working
electrode. Still another embodiment includes having a digoxin or
digoxigenin redox mediator react with antidigoxin in a matrix near
or immobilized on a working electrode.
Preferred non-leachable redox mediators are redox polymers, such as
polymeric transition metal compounds or complexes. Typically, the
polymers used to form a redox polymer have nitrogen-containing
heterocycles, such as pyridine, imidazole, or derivatives thereof
for binding as ligands to the redox species. Suitable polymers for
complexation with redox species, such as the transition metal
complexes, described above, include, for example, polymers and
copolymers of poly(1-vinyl imidazole) (referred to as "PVI") and
poly(4-vinyl pyridine) (referred to as "PVP"), as well as polymers
and copolymers of poly(acrylic acid) or polyacrylamide that have
been modified by the addition of pendant nitrogen-containing
heterocycles, such as pyridine and imidazole. Modification of
poly(acrylic acid) may be performed by reaction of at least a
portion of the carboxylic acid functionalities with an
aminoalkylpyridine or aminoalkylimidazole, such as
4-ethylaminopyridine, to form amides. Suitable copolymer
substituents of PVI, PVP, and poly(acrylic acid) include
acrylonitrile, acrylamide, acrylhydrazide, and substituted or
quaternized 1-vinyl imidazole. The copolymers can be random or
block copolymers.
The transition metal complexes of non-leachable redox polymers are
typically covalently or coordinatively bound with the
nitrogen-containing heterocycles (e.g., imidazole and/or pyridine
rings) of the polymer. The transition metal complexes may have
vinyl functional groups through which the complexes can be
co-polymerized. Suitable vinyl functional groups include, for
example, vinylic heterocycles, amides, nitriles, carboxylic acids,
sulfonic acids, or other polar vinylic compounds. An example of a
redox polymer of this type is poly(vinyl ferrocene) or a derivative
of poly(vinyl ferrocene) functionalized to increase swelling of the
redox polymer in water.
Another type of redox polymer contains an ionically-bound redox
species, by forming multiple ion-bridges. Typically, this type of
mediator includes a charged polymer coupled to an oppositely
charged redox species. Examples of this type of redox polymer
include a negatively charged polymer such as Nafion.RTM. (DuPont)
coupled to multiple positively charged redox species such as an
osmium or ruthenium polypyridyl cation. Another example of an
ionically-bound mediator is a positively charged polymer such as
quaternized poly(4-vinyl pyridine) or poly(1-vinyl imidazole)
coupled to a negatively charged redox species such as ferricyanide
or ferrocyanide. The preferred ionically-bound redox species is a
multiply charged, often polyanionic, redox species bound within an
oppositely charged polymer.
Another suitable redox polymer includes a redox species
coordinatively bound to a polymer. For example, the mediator may be
formed by coordination of an osmium, ruthenium, or cobalt
2,2'-bipyridyl complex to poly(1-vinyl imidazole) or poly(4-vinyl
pyridine) or by co-polymerization of, for example, a
4-vinyl-2,2'-bipyridyl osmium, ruthenium, or cobalt complex with
1-vinyl imidazole or 4-vinyl pyridine.
Typically, the ratio of osmium or ruthenium transition metal
complexes to imidazole and/or pyridine groups of the non-leachable
redox polymers ranges from 1:20 to 1:1, preferably from 1:15 to
1:2, and more preferably from 1:10 to 1:4. Generally, the redox
potentials depend, at least in part, on the polymer with the order
of redox potentials being poly(acrylic acid)<PVI<PVP.
A variety of methods may be used to immobilize a redox polymer on
an electrode surface. One method is adsorptive immobilization. This
method is particularly useful for redox polymers with relatively
high molecular weights. The molecular weight of a polymer may be
increased, for example, by cross-linking. The polymer of the redox
polymer may contain functional groups, such as, for example,
hydrazide, amine, alcohol, heterocyclic nitrogen, vinyl, allyl, and
carboxylic acid groups, that can be crosslinked using a
crosslinking agent. These functional groups may be provided on the
polymer or one or more of the copolymers. Alternatively or
additionally, the functional groups may be added by a reaction,
such as, for example, quaternization. One example is the
quaternization of PVP with bromoethylamine groups.
Suitable cross-linking agents include, for example, molecules
having two or more epoxide (e.g., poly(ethylene glycol) diglycidyl
ether (PEGDGE)), aldehyde, aziridine, alkyl halide, and azide
functional groups or combinations thereof. When a polymer has
multiple acrylate functions, it can be crosslinked with a di- or
polythiol; when it has multiple thiol functions it can be
crosslinked with a di- or polyacrylate. Other examples of
cross-linking agents include compounds that activate carboxylic
acid or other acid functional groups for condensation with amines
or other nitrogen compounds. These cross-linking agents include
carbodiimides or compounds with active N-hydroxysuccinimide or
imidate functional groups. Yet other examples of cross-linking
agents are quinones (e.g., tetrachlorobenzoquinone and
tetracyanoquinodimethane) and cyanuric chloride. Other
cross-linking agents may also be used. In some embodiments, an
additional cross-linking agent is not required. Further discussion
and examples of cross-linking and cross-linking agents are found in
U.S. Pat. Nos. 5,262,035; 5,262,305; 5,320,725; 5,264,104;
5,264,105; 5,356,786; and 5,593,852, herein incorporated by
reference.
In another embodiment, the redox polymer is immobilized by the
functionalization of the electrode surface and then the chemical
bonding, often covalently, of the redox polymer to the functional
groups on the electrode surface. One example of this type of
immobilization begins with a poly(4-vinyl pyridine). The polymer's
pyridine rings are, in part, complexed with a reducible/oxidizable
species, such as [Os(bpy).sub.2Cl].sup.+/2+ where bpy is
2,2'-bipyridine. Part of the pyridine rings are quaternized by
reaction with 2-bromoethylamine. The polymer is then crosslinked,
for example, using a diepoxide, such as poly(ethylene glycol)
diglycidyl ether.
Carbon surfaces can be modified for attachment of a redox polymer,
for example, by electroreduction of a diazonium salt. As an
illustration, reduction of a diazonium salt formed upon
diazotization of p-aminobenzoic acid modifies a carbon surface with
phenylcarboxylic acid functional groups. These functional groups
can be activated by a carbodiimide, such as
1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide hydrochloride (EDC).
The activated functional groups are bound with an
amine-functionalized redox couple, such as, for example, the
quaternized osmium-containing redox polymer described above or
2-aminoethylferrocene, to form the redox couple.
Similarly, gold and other metal surfaces can be functionalized by,
for example, an amine, such as cystamine, or by a carboxylic acid,
such as thioctic acid. A redox couple, such as, for example,
[Os(bpy).sub.2(pyridine-4-carboxylate)Cl].sup.0/+, is activated by
1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide hydrochloride (EDC)
to form a reactive O-acylisourea which reacts with the gold-bound
amine to form an amide. The carboxylic acid functional group of
thioctic acid can be activated with EDC to bind a polymer or
protein amine to form an amide.
When the enzyme used is PQQ glucose dehydrogenase or glucose
oxidase, the preferred non-leachable redox mediators have a redox
potential between about -300 mV to about +400 mV versus the
standard calomel electrode (SCE). The most preferred non-leachable
redox mediators have osmium redox centers and a redox potential
more negative than +100 mV versus SCE, more preferably the redox
potential is more negative than 0 mV versus SCE, and most
preferably is near -150 mV versus SCE.
In at least some instances, the redox mediators of the sensors are
air-oxidizable. This means that the redox mediator is oxidized by
air, preferably, so that at least 90% of the mediator is in an
oxidized state prior to introduction of sample into the sensor.
Air-oxidizable redox mediators include osmium cations complexed
with two mono-, di-, or polyalkoxy-2,2'-bipyridine or mono-, di-,
or polyalkoxy-1,10-phenanthroline ligands, the two ligands not
necessarily being the same, and further complexed with polymers or
other ligands having pyridine and imidazole functional groups. In
particular, Os[4,4'-dimethoxy-2,2'-bipyridine].sub.2Cl.sup.+/+2
complexed with poly(4-vinyl pyridine) or poly(1-vinyl imidazole)
attains approximately 90% or more oxidation in air. The air
oxidation of the redox mediator may take place while the redox
mediator is a solid, such as, for example, when it is coated on the
sensor in a dry state and stored. Alternatively, the air oxidation
of the redox mediator may take place while the redox mediator is in
solution, such as, for example, prior to the solution being applied
onto the sensor and dried. In the case in which the redox mediator
is air oxidized in solution, the solution containing the redox
mediator may be kept in storage for a period of time sufficient to
air oxidize the mediator prior to use of the solution in the
manufacturing process.
Second Electron Transfer Agent
In a preferred embodiment of the invention, the sensor includes a
redox mediator and a second electron transfer agent which is
capable of transferring electrons to or from the redox mediator and
the analyte. The second electron transfer agent may be diffusible
or may be non-leachable (e.g., entrapped in or coordinatively,
covalently, or ionically bound to the redox polymer). One example
of a suitable second electron transfer agent is an enzyme which
catalyzes a reaction of the analyte. For example, a glucose oxidase
or glucose dehydrogenase, such as pyrroloquinoline quinone glucose
dehydrogenase (PQQ), is used when the analyte is glucose. A lactate
oxidase fills this role when the analyte is lactate. Other enzymes
can be used for other analytes. These enzymes catalyze the
electrolysis of an analyte by transferring electrons between the
analyte and the electrode via the redox mediator. In some
embodiments, the second electron transfer agent is non-leachable,
and more preferably immobilized on the working electrode, to
prevent unwanted leaching of the agent into the sample. This is
accomplished, for example, by cross-linking the non-leachable
second electron transfer agent with the non-leachable redox
mediator, thereby providing a sensing layer with non-leachable
components on the working electrode. In other embodiments, the
second electron transfer agent is diffusible (and may be disposed
on any surface of the sample chamber or placed in the sample).
Counter Electrode
Counter electrode 24, as illustrated in FIGS. 1-4, may be
constructed in a manner similar to working electrode 22. Counter
electrode 24 may also be a counter/reference electrode.
Alternatively, a separate reference electrode may be provided in
contact with the sample chamber. Suitable materials for the
counter/reference or reference electrode include Ag/AgCl or Ag/AgBr
printed on a non-conducting base material or silver chloride on a
silver metal base. The same materials and methods may be used to
make the counter electrode as are available for constructing the
working electrode 22, although different materials and methods may
also be used. A tab 25 may be provided on the electrode for
convenient connection to the external electronics (not shown), such
as a coulometer, potentiostat, or other measuring device.
Electrode Configuration
In one embodiment of the invention, working electrode 22 and
counter electrode 24 are disposed opposite to and facing each other
to form a facing electrode pair as depicted in FIGS. 1 and 3. In
this preferred configuration, the sample chamber 26 is typically
disposed between the two electrodes. For this facing electrode
configuration, it is preferred that the electrodes are separated by
a distance of no more than about 0.2 mm (i.e., at least one portion
of the working electrode is separated from one portion of the
counter electrode by no more than 200 .mu.m), preferably no more
than 100 .mu.m, and most preferably no more than 50 .mu.m.
The electrodes need not be directly opposing each other; they may
be slightly offset. Furthermore, the two electrodes need not be the
same size. Preferably, the counter electrode 24 is at least as
large as the working surface of the working electrode 22. The
counter electrode 22 can also be formed with tines in a comb shape.
Other configurations of both the counter electrode and working
electrode are within the scope of the invention. However, for this
particular embodiment, the separation distance between at least one
portion of the working electrode and some portion of the counter
electrode preferably does not exceed the limits specified
hereinabove.
FIGS. 11A, 11B, and 11C illustrate different embodiments of pairs
of facing electrodes 22, 24, as described above. A region 21 of
overlap between the two electrodes 22, 24 typically corresponds to
the measurement zone in which the sample will be interrogated. Each
of the electrodes 22, 24 is a conducting surface and acts as a
plate of a capacitor. The measurement zone between the electrodes
22, 24 acts as a dielectric layer between the plates. Thus, there
is a capacitance between the two electrodes 22, 24. This
capacitance is a function of the size of the overlapping electrodes
22, 24, the separation between the electrodes 22, 24, and the
dielectric constant of the material between the electrodes 22, 24.
Thus, if the size of the region 21 of the overlapping electrodes
22, 24 and the dielectric constant of the material between the
electrodes (e.g., air or a sorbent material) are known, then the
separation between the electrodes can be calculated to determine
the volume of the measurement zone.
FIG. 11A illustrates one embodiment of the invention in which the
electrodes 22, 24 are positioned in a facing arrangement. For the
capacitance to be uniform among similarly constructed analyte
sensors having this particular sensor configuration, the
registration (i.e., the positioning of the two electrodes relative
to one another) should be uniform. If the position of either of the
electrodes is shifted in the x-y plane from the position shown in
FIG. 11A, the size of the overlap, and therefore, of the
capacitance, will change. The same principle holds for the volume
of the measurement zone.
FIGS. 11B and 11C illustrate other embodiments of the invention
with electrodes 22, 24 in a facing arrangement. In these particular
arrangements, the position of either of the electrodes may be
shifted, by at least some minimum distance, in the x-y plane
relative to the other electrode without a change in the capacitance
or the volume of the measurement zone. In these electrode
arrangements, each electrode 22, 24 includes an arm 122, 124,
respectively, which overlaps with the corresponding arm of the
other electrode. The two arms 122, 124 are not parallel to each
other (such as illustrated in FIG. 11A); rather, the arms 122, 124
are disposed at an angle 123, which is greater than zero, relative
to each other. In addition, the two arms 122, 124 extend beyond the
region 21 of overlap (i.e., each arm has extra length corresponding
to the difference between the length of the arm 222, 224,
respectively, and the width 121 of the overlap 21). With these
electrode arrangements, there can be a certain amount of allowed
imprecision in the registration of the electrodes 22, 24 which does
not change the capacitance of the electrode arrangement. A desired
amount of allowed imprecision in the registration can be designed
into the electrode arrangement by varying the angle 123 at which
the arms 122, 124 overlap and the size of the extra length of each
arm 122, 124 relative to the width 121 of the region 21 of overlap.
Typically, the closer that the arms 122, 124 are to being
perpendicular (i.e., angle 123 is 90.degree.), the greater the
allowed imprecision. Also, the greater the extra length of each arm
122, 124 (which may both be the same length or different lengths)
relative to the width 121 of the region 21 of overlap, the greater
the allowed imprecision. Conversely, the greater the amount of
allowed imprecision, the larger the size of the electrodes (for a
given electrode width, thickness, and angle 123 of intersection
with the other electrode). Thus, the minimum distance that one
electrode can be shifted relative to the other is balanced against
the amount of material needed for the electrodes. Typically, the
angle 123 of intersection ranges from 5 to 90 degrees, preferably,
30 to 90 degrees, and more preferably 60 to 90 degrees. Typically,
the ratio of the extra length of an arm 122, 124 (corresponding to
the difference between the arm length 222, 224 and the width 121 of
the region 21 of overlap) versus the width 121 of the region 21 of
overlap ranges from 0.1:1 to 50:1, preferably 1:1 to 15:1, and more
preferably 4:1 to 10:1.
In another embodiment of the invention, the two electrodes 22, 24
are coplanar as shown in FIG. 2. In this case, the sample chamber
26 is in contact with both electrodes and is bounded on the side
opposite the electrodes by a non-conducting inert base 30. Suitable
materials for the inert base include non-conducting materials such
as polyester.
Other configurations of the inventive sensors are also possible.
For example, the two electrodes may be formed on surfaces that make
an angle to each other. One such configuration would have the
electrodes on surfaces that form a right angle. Another possible
configuration has the electrodes on a curved surface such as the
interior of a tube. The working and counter electrodes may be
arranged so that they face each other from opposite sides of the
tube. This is another example of a facing electrode pair.
Alternatively, the electrodes may be placed near each other on the
tube wall (e.g., one on top of the other or side-by-side).
In any configuration, the two electrodes must be configured so that
they do not make direct electrical contact with each other, to
prevent shorting of the electrochemical sensor. This may be
difficult to avoid when the facing electrodes are separated, over
the average, by no more than about 100 .mu.m.
A spacer 28 can be used to keep the electrodes apart when the
electrodes face each other as depicted in FIGS. 1 and 3. The spacer
is typically constructed from an inert non-conducting material such
as pressure-sensitive adhesive, polyester, Mylar.TM., Kevlar.TM. or
any other strong, thin polymer film, or, alternatively, a thin
polymer film such as a Teflon.TM. film, chosen for its chemical
inertness. In addition to preventing contact between the
electrodes, the spacer 28 often functions as a portion of the
boundary for the sample chamber 26 as shown in FIGS. 1-4. Other
spacers include layers of adhesive and double-sided adhesive tape
(e.g., a carrier film with adhesive on opposing sides of the
film).
Sample Chamber
The sample chamber 26 is typically defined by a combination of the
electrodes 22, 24, an inert base 30, and a spacer 28 as shown in
FIGS. 1-4. A measurement zone is contained within this sample
chamber and is the region of the sample chamber that contains only
that portion of the sample that is interrogated during the analyte
assay. In the embodiment of the invention illustrated in FIGS. 1
and 2, sample chamber 26 is the space between the two electrodes
22, 24 and/or the inert base 30. In this embodiment, the sample
chamber has a volume that is preferably no more than about 1 .mu.L,
more preferably no more than about 0.5 .mu.L, and most preferably
no more than about 0.25 .mu.L. In the embodiment of the invention
depicted in FIGS. 1 and 2, the measurement zone has a volume that
is approximately equal to the volume of the sample chamber. In a
preferred embodiment the measurement zone includes 80% of the
sample chamber, 90% in a more preferred embodiment, and about 100%
in a most preferred embodiment.
In another embodiment of the invention, shown in FIG. 3, sample
chamber 26 includes much more space than the region proximate
electrodes 22, 24. This configuration makes it possible to provide
multiple electrodes in contact with one or more sample chambers, as
shown in FIG. 5. In this embodiment, sample chamber 26 is
preferably sized to contain a volume of no more than about 1 .mu.L,
more preferably no more than about 0.5 .mu.L, and most preferably
no more than about 0.25 .mu.L. The measurement zone (i.e., the
region containing the volume of sample to be interrogated) is
generally sized to contain a volume of sample of no more than about
1 .mu.L, preferably no more than about 0.5 .mu.L, more preferably
no more than about 0.25 .mu.L, and most preferably no more than
about 0.1 .mu.L. One particularly useful configuration of this
embodiment positions working electrode 22 and counter electrode 24
facing each other, as shown in FIG. 3. In this embodiment, the
measurement zone, corresponding to the region containing the
portion of the sample which will be interrogated, is the portion of
sample chamber 26 bounded by the working surface of the working
electrode and disposed between the two facing electrodes.
In both of the embodiments discussed above, the thickness of the
sample chamber and of the measurement zone correspond typically to
the thickness of spacer 28 (e.g., the distance between the
electrodes in FIGS. 1 and 3, or the distance between the electrodes
and the inert base in FIG. 2). The spacer may be, for example, an
adhesive or double-sided adhesive tape or film. Preferably, this
thickness is small to promote rapid electrolysis of the analyte, as
more of the sample will be in contact with the electrode surface
for a given sample volume. In addition, a thin sample chamber helps
to reduce errors from diffusion of analyte into the measurement
zone from other portions of the sample chamber during the analyte
assay, because diffusion time is long relative to the measurement
time. Typically, the thickness of the sample chamber is no more
than about 0.2 mm. Preferably, the thickness of the sample chamber
is no more than about 0.1 mm and, more preferably, the thickness of
the sample chamber is about 0.05 mm or less.
The sample chamber may be formed by other methods. Exemplary
methods include embossing, indenting, or otherwise forming a recess
in a substrate within which either the working electrode 22 or
counter electrode 24 is formed. FIGS. 12A and 12B illustrate one
embodiment of this structure. First, a conducting layer 100 is
formed on an inert non-conducting base material 102. As described
above, the conducting layer 100 can include gold, carbon, platinum,
ruthenium dioxide, palladium, or other non-corroding materials. The
inert non-conducting base material 102 can be made using a
polyester, other polymers, or other non-conducting, deformable
materials. A recess 104 is then formed in a region of the
non-conducting base material 102 so that at least a portion of the
conducting layer 100 is included in the recess 104. The recess 104
may be formed using a variety of techniques including indenting,
deforming, or otherwise pushing in the base material 102. One
additional exemplary method for forming the recess includes
embossing the base material 102. For example, the base material 102
may be brought into contact with an embossing roll or stamp having
raised portions, such as punch members or channels, to form the
recess 104. In some embodiments, the base material 102 may be
heated to soften the material.
The recess 104 may be circular, oval, rectangular, or any other
regular or irregular shape. Alternatively, the recess 104 may be
formed as a channel which extends along a portion of the base
material 102. The conducting layer 100 may extend along the entire
channel or only a portion of the channel. The measurement zone may
be restricted to a particular region within the channel by, for
example, depositing the sensing layer 32 on only that portion of
the conducting layer 100 within the particular region of the
channel. Alternatively, the measurement zone may be defined by
placing a second electrode 107 over only the desired region of the
first electrode 105.
At least a portion, and in some cases, all, of the conducting layer
100 is situated in the recess 104. This portion of the conducting
layer 100 may act as a first electrode 105 (a counter electrode or
a working electrode). If the conducting layer 100 forms the working
electrode, then a sensing layer 32 may be formed over a portion of
the conducting layer 100 by depositing a non-leachable redox
mediator and/or second electron transfer agent in the recess 104,
as shown in FIG. 12B. If a diffusible redox mediator or second
electron transfer agent is used, then the diffusible material may
be disposed on any surface in the sample chamber or in the
sample.
A second electrode 107 is then formed by depositing a second
conducting layer on a second base material 106. This second
electrode 107 is then positioned over the first electrode 105 in a
facing arrangement. Although not illustrated, if the redox mediator
is non-leachable it will be understood that if the first electrode
105 were to function as a counter electrode, then the sensing layer
32 would be deposited on the second electrode 107 which would then
function as the working electrode. If the redox mediator is
diffusible, however, the redox mediator may be disposed on any
surface of the sample chamber or may be placed in the sample.
In one embodiment, the second base material 106 rests on a portion
of the first base material 102 and/or the conducting layer 100
which is not depressed, so that the second electrode 107 extends
into the recess. In another embodiment, there is a spacer (not
shown) between the first and second base materials 102, 106. In
this embodiment, the second electrode 107 may or may not extend
into the recess. In any case, the first and second electrodes 105,
107 do not make contact, otherwise the two electrodes would be
shorted.
The depth of the recess 104 and the volume of the conductive layer
100, sensing layer 32, and the portion, if any, of the second
electrode 107 in the recess 104 determines the volume of the
measurement zone. Thus, the predictability of the volume of the
measurement zone relies on the extent to which the formation of the
recess 104 is uniform.
In addition to the conducting layer 100, a sorbent layer 103,
described in detail below, may be deposited on the base material
102 prior to forming the recess 104, as shown in FIG. 14A. The
sorbent material 103 may be indented, embossed, or otherwise
deformed with the conducting layer 100 and base material 102, as
shown in FIG. 14B. Alternatively, the sorbent material 103 may be
deposited after the conducting layer 100 and base material 102 are
indented, embossed, or otherwise deformed to make the recess
104.
In another exemplary method for forming the analyte sensor, a
recess 114 is formed in a first base material 112, as shown in
FIGS. 13A and 13B. The recess may be formed by indenting,
embossing, etching (e.g., using photolithographic methods or laser
removal of a portion of the base material), or otherwise deforming
or removing a portion of the base material 112. Then a first
conducting layer 110 is formed in the recess 114. Any of the
conductive materials discussed above may be used. A preferred
material is a conductive ink, such as a conductive carbon ink
available, for example, from Ercon, Inc. (Wareham, Mass.). The
conductive ink typically contains metal or carbon dissolved or
dispersed in a solvent or dispersant. When the solvent or
dispersant is removed, the metal or carbon forms a conductive layer
110 that can then be used as a first electrode 115. A second
electrode 117 can be formed on a second base material 116 and
positioned over the recess 114, as described above. In embodiments
having a non-leachable redox mediator, a sensing layer 32 is formed
on the first electrode 115 to form a working electrode, as shown in
FIG. 13B. In other embodiments having a non-leachable redox
mediator, the sensing layer 32 may be formed on the second
electrode 117 to form a working electrode. Alternatively, if a
diffusible redox mediator is used, then the working electrode need
not include the sensing layer disposed thereon. In fact, no sensing
layer is required because the redox mediator may be placed in the
sample and likewise for a diffusible second electron transfer
agent, if one is present. Any diffusible components may be
independently disposed on any surface of the sample chamber or
placed in the sample. Furthermore, a sorbent material (not shown)
may be formed within the recess, for example, on the first
electrode 115.
A binder, such as a polyurethane resin, cellulose derivative,
elastomer (e.g., silicones, polymeric dienes, or
acrylonitrile-butadiene-styrene (ABS) resins), highly fluorinated
polymer, or the like, may also be included in the conductive ink.
Curing the binder may increase the conductivity of the conductive
layer 110, however, curing is not necessary. The method of curing
the binder may depend on the nature of the particular binder that
is used. Some binders are cured by heat and/or ultraviolet
light.
These structures allow for the formation of electrochemical sensors
in which the volume of the measurement zone depends, at least in
part, on the accuracy and reproducibility of the recess 104.
Embossing, laser etching, photolithographic etching and other
methods can be used to make reproducible recesses 104, even on the
scale of 200 .mu.m or less.
Sorbent Material
The sample chamber may be empty before the sample is placed in the
chamber. Alternatively, the sample chamber may include a sorbent
material 34 to sorb and hold a fluid sample during the measurement
process. Suitable sorbent materials include polyester, nylon,
cellulose, and cellulose derivatives such as nitrocellulose. The
sorbent material facilitates the uptake of small volume samples by
a wicking action which may complement or, preferably, replace any
capillary action of the sample chamber. In addition or
alternatively, a portion or the entirety of the wall of the sample
chamber may be covered by a surfactant, such as, for example, Zonyl
FSO.
In some embodiments, the sorbent material is deposited using a
liquid or slurry in which the sorbent material is dissolved or
dispersed. The solvent or dispersant in the liquid or slurry may
then be driven off by heating or evaporation processes. Suitable
sorbent materials include, for example, cellulose or nylon powders
dissolved or dispersed in a suitable solvent or dispersant, such as
water. The particular solvent or dispersant should also be
compatible with the material of the working electrode 22 (e.g., the
solvent or dispersant should not dissolve the electrode).
One of the most important functions of the sorbent material is to
reduce the volume of fluid needed to fill the sample chamber and
corresponding measurement zone of the sensor. The actual volume of
sample within the measurement zone is partially determined by the
amount of void space within the sorbent material. Typically,
suitable sorbents consist of about 5% to about 50% void space.
Preferably, the sorbent material consists of about 10% to about 25%
void space.
The displacement of fluid by the sorbent material is advantageous.
By addition of a sorbent, less sample is needed to fill sample
chamber 26. This reduces the volume of sample that is required to
obtain a measurement and also reduces the time required to
electrolyze the sample.
The sorbent material 34 may include a tab 33 which is made of the
same material as the sorbent and which extends from the sensor, or
from an opening in the sensor, so that a sample may be brought into
contact with tab 33, sorbed by the tab, and conveyed into the
sample chamber 26 by the wicking action of the sorbent material 34.
This provides a preferred method for directing the sample into the
sample chamber 26. For example, the sensor may be brought into
contact with a region of an animal (including human) that has been
pierced with a lancet to draw blood. The blood is brought in
contact with tab 33 and drawn into sample chamber 26 by the wicking
action of the sorbent 34. The direct transfer of the sample to the
sensor is especially important when the sample is very small, such
as when the lancet is used to pierce a portion of the animal that
is not heavily supplied with near-surface capillary vessels and
furnishes a blood sample volume of 1 .mu.L or less.
Methods other than the wicking action of a sorbent may be used to
transport the sample into the sample chamber or measurement zone.
Examples of such methods for transport include the application of
pressure on a sample to push it into the sample chamber, the
creation of a vacuum by a pump or other vacuum-producing method in
the sample chamber to pull the sample into the chamber, capillary
action due to interfacial tension of the sample with the walls of a
thin sample chamber, as well as the wicking action of a sorbent
material.
The sensor can also be used in conjunction with a flowing sample
stream. In this configuration, the sample stream is made to flow
through a sample chamber. The flow is stopped periodically and the
concentration of the analyte is determined by an electrochemical
method, such as coulometry. After the measurement, the flow is
resumed, thereby removing the sample from the sensor.
Alternatively, sample may flow through the chamber at a very slow
rate, such that all of the analyte is electrolyzed in transit,
yielding a current dependent only upon analyte concentration and
flow rate.
Other filler materials may be used to fill the measurement zone and
reduce the sample volume. For example, glass beads can be deposited
in the measurement zone to occupy space. Preferably, these filler
materials are hydrophilic so that the body fluid can easily flow
into the measurement zone. In some cases, such as glass beads with
a high surface area, these filler materials may also wick the body
fluid into the measurement zone due to their high surface area and
hydrophilicity.
The entire sensor assembly is held firmly together to ensure that
the sample remains in contact with the electrodes and that the
sample chamber and measurement zone maintain the same volume. This
is an important consideration in the coulometric analysis of a
sample, where measurement of a defined sample volume is needed. One
method of holding the sensor together is depicted in FIGS. 1 and 2.
Two plates 38 are provided at opposite ends of the sensor. These
plates are typically constructed of non-conducting materials such
as plastics. The plates are designed so that they can be held
together with the sensor between the two plates. Suitable holding
devices include adhesives, clamps, nuts and bolts, screws, and the
like.
Alternative Sensor Designs
FIGS. 18A to 18C illustrate one alternative sensor design for
formation of thin film sensors. The sensor includes a first
substrate 500 upon which a working electrode 502 is formed. The
working electrode 502 includes a contact region 503 for connection
with external electronics. A spacer 504 (FIG. 18B), such as, for
example, a layer of adhesive or a double-sided tape defines a
channel 506 to produce a sample chamber for the sensor. Two counter
(or counter/reference) electrodes 510, 512 are formed on a second
substrate 508, as shown in FIG. 18C (inverted with respect to FIGS.
18A and 18B to show the electrode side up). This multiple counter
electrode arrangement may provide a fill indicator function, as
described below. Each counter electrode 510, 512 has a contact
region 511, 513 for connection with external electronics. The
second substrate 508 is inverted and placed over the first
substrate 500, with the spacer 504 between, so that the working
electrode 502 and the two counter electrodes 510, 512 are facing in
the region of the channel 506.
In some instances, the counter electrode 510 nearest an entrance
514 of the channel 506 has a surface area within the sample chamber
that is at least two times larger than the other counter electrode
512, and may be at least five or ten times larger. The
non-leachable or diffusible redox mediator and/or second electron
transfer agent can be provided on either the first or second
substrates 500, 508 in a region corresponding to the channel 506,
as described above.
The working electrode and counter electrodes can be formed to cover
the entire channel region (except for a small space between the two
counter electrodes). In this embodiment, the sample chamber and
measurement zone are effectively the same and have the same volume.
In other embodiments, the measurement zone has, for example, 80% or
90% of the volume of the sample chamber. It will be understood that
similar sensors could be made using one counter electrode or three
or more counter electrodes. It will also be understood that
multiple working electrodes may also be provided on the sensor.
One example of a method for making the thin film sensors is
described with respect to the sensor arrangement displayed in FIGS.
18A to 18C and can be used to make a variety of other sensor
arrangements, including those described before. A substrate, such
as a plastic substrate, is provided. The substrate can be an
individual sheet or a continuous roll on a web. This substrate can
be used to make a single sensor or to make multiple sensors. The
multiple sensors can be formed on a substrate 1000 as working
electrodes 1010 and counter electrode(s) 1020. In some embodiments,
the substrate can be scored and folded to bring the working
electrodes 1010 and counter electrodes 1020 together to form the
sensor. In some embodiments, as illustrated in FIG. 31A, the
individual working electrodes 1010 (and, in a separate section, the
counter electrode(s) 1020) can be formed next to each other on the
substrate 1000, to reduce waste material, as illustrated in FIG.
31A. In other embodiments, the individual working electrodes 1010
(and, in a separate section, the counter electrode(s) 1020) can be
spaced apart, as illustrated in FIG. 31B. The remainder of the
process is described for the manufacture of multiple sensors, but
can be readily modified to form individual sensors.
Carbon or other electrode material (e.g., metal, such as gold or
platinum) is formed on the substrate to provide a working electrode
for each sensor. The carbon or other electrode material can be
deposited by a variety of methods including printing a carbon or
metal ink, vapor deposition, and other methods.
Optionally, a non-conductive material, such as a non-conductive
ink, can be formed adjacent the working electrode to provide a
planar surface along the path of travel of the sample fluid. The
non-conductive material is suitable for creating a smooth surface
to facilitate filling by capillary action and/or for reducing the
likelihood that air bubbles will become entrapped near the working
electrode. This non-conductive material can be colored or colorless
and may be formed on the substrate by printing or other techniques.
The non-conductive material may be deposited prior to or subsequent
to the formation of the working electrode.
The counter electrode or counter electrodes are formed on the
substrate. The counter electrode(s) are formed by depositing carbon
or other electrode material onto the substrate. In one embodiment,
the material of the counter electrode(s) is a Ag/AgCl ink. The
material of the counter electrode(s) may be deposited by a variety
of methods including printing or vapor deposition. In some
embodiments, the counter electrodes are formed using different
materials and/or one electrode is a counter or counter/reference
electrode and the other electrode is a reference or
counter/reference electrode. In one embodiment, the working
electrodes are formed on a first half of a polymer sheet or web and
the counter electrodes are formed on a second half of the polymer
sheet or web so that the sheet or web can be folded to superimpose
the working and counter electrodes in a facing arrangement.
A second non-conductive material may be deposited adjacent and/or
between the counter electrode(s) to provide a planar surface along
the path of travel of the sample fluid. This may be particularly
desirable in the region between the counter electrodes that will be
part of the sample chamber to planarize the surface of the sample
chamber. The non-conductive material is suitable for creating a
smooth surface to faciliate filling by capillary action and/or for
reducing the likelihood that air bubbles will become entrapped
between or near the counter electrode(s). This non-conductive
material can be colored or colorless and may be formed on the by
printing or other techniques. The non-conductive material may be
deposited prior to or subsequent to the formation of the counter
electrode(s).
An adhesive spacer is formed over at least one of the
substrate/working electrode and substrate/counter electrode(s). The
adhesive spacer may be a single layer of adhesive or a double-sided
adhesive tape (e.g., a polymer carrier film with adhesive disposed
on opposing surfaces). To form the channel, the spacer, optionally
provided with one or more release liners, may be cut (e.g.,
die-cut) to remove the portion of the adhesive corresponding to the
channel prior to disposing the spacer on the substrate.
Alternatively, the adhesive may be printed or otherwise disposed on
the substrate according to a pattern which defines the channel
region. The thickness of the spacer typically determines the
spacing between the working and counter electrodes. When the
uniformity of this spacing among sensors is necessary (e.g., for
coulometric measurements), uniformity in the thickness of the
spacer is important. Preferably, the thickness does not vary more
than .+-.5% over the individual sensor and/or among individual
sensors in a batch.
The non-leachable or diffusible redox mediator and/or second
electron transfer agent are disposed onto the substrate in at least
the sample chamber region. If either or both of these components is
non-leachable, that component or components must be disposed on the
working electrode. If either or both of these components is
diffusible, that component or components can be disposed on any
surface of the substrate in the channel region. The redox mediator
and/or second electrode transfer agent can be disposed
independently or together on the substrate prior to or after
disposition of the spacer. The redox mediator and/or second
electrode transfer agent may be disposed by a variety of methods
including, for example, screen printing, ink jet printing,
spraying, painting, striping along a row or column of aligned
and/or adjacent electrodes, and the like. Other components may be
deposited separately or with the redox mediator and/or second
electrode transfer agent including, for example, surfactants,
polymers, polymer films, preservatives, binders, buffers, and
cross-linkers.
After disposing the spacer, redox mediator, and second electron
transfer agent, the substrate can be folded to form the sensor. The
faces of the substrate are joined by the adhesive of the spacer.
After bringing the faces together, the sensor can be cut out using
a variety of methods including, for example, die cutting, slitting,
or otherwise cutting away the excess substrate material and
separating the individual sensors. In some embodiments, a
combination of methods may be used. For example, some features may
be die cut, while the remainder of the sensor is cut by slitting.
As another alternative, the sensor components (e.g., the components
illustrated in FIGS. 18A and 18C) may first be cut out of the
substrates and then brought together to form the sensor by
adhesively joining the two components using the spacer
adhesive.
The embodiment of a sensor illustrated in FIGS. 18A to 18C is an
example of a tip-fill sensor. An alternative sensor construction is
illustrated in FIGS. 19A to 19C. This is a side-fill sensor. FIG.
19A illustrates a first substrate 520 with a working electrode 522.
FIG. 19B illustrates a spacer 524 defining a channel 526. FIG. 19C
(inverted with respect to FIGS. 19A and 19B to illustrate the
electrodes) illustrates a second substrate 528 with three counter
(or counter/reference) electrodes 530, 532, 534.
This sensor can be manufactured as described above. The symmetric
disposition of the counter electrodes allow the sensor to be filled
from either the left or right side for convenience of left-handed
and right-handed people. It will be understood, however, that
similar sensor arrangements can be formed using one, two, or four
or more counter electrode(s) and/or two or more working electrodes.
The scalloped regions 536, 538 may be formed, for example, by die
cutting and may, at least in some instances, be precisely
controlled to provide a reproducible channel length. As an
alternative arrangement, the sides of the sensor may be straight to
allow the sensor to be cut out from the remainder of the substrate
and/or from other sensors by slitting the substrate in parallel
directions using, for example, a gang arbor blade system. As
illustrated in FIGS. 19A, 19B, and 19C, the edges of the sensor can
define edges of the sample chamber and/or measurement zone. By
accurately controlling the distance between cuts, variability in
sample chamber volume can often be reduced. In some instances,
these cuts are preferably parallel to each other, as parallel cuts
may be the easiest to make reproducibly.
FIGS. 20A, 20B, and 20C illustrate another example of a
side-filling sensor arrangement. FIG. 20A illustrates a first
substrate 540 with a working electrode 542. FIG. 20B illustrates a
spacer 544 defining a channel 546. FIG. 20C (inverted with respect
to FIGS. 20A and 2013) illustrates a second substrate 548 with
three counter (or counter/reference) electrodes 550, 552, 554.
FIGS. 21A, 21B, and 21C illustrate another example of a tip-filling
sensor arrangement. FIG. 21A illustrates a first substrate 560 with
a working electrode 562. FIG. 21B illustrates a spacer 564 defining
a channel 566. FIG. 21C (inverted with respect to FIGS. 21A and
21B) illustrates a second thin film substrate 568 with two counter
(or counter/reference) electrodes 570, 572. A vent hole 574
(indicated as a shaded region in FIG. 21C) is provided through the
second substrate. In the illustrated embodiment, this vent hole 574
is made through only the substrate 568 that carries the counter
electrode(s) and, optionally, the spacer 564. In this embodiment,
the vent hole may be formed by, for example, die cutting a portion
of the substrate. This die cut may remove a portion of at least one
counter electrode, but a sufficient amount of the counter electrode
should remain for contact with the sample in the channel and for
electrical connection to a contact at the other end of the sensor.
In another embodiment, the vent hole 574 may be made through all of
the layers or through the first substrate and not the second
substrate.
Another embodiment is illustrated in FIGS. 22A, 22B, and 22C, with
a different shape. This sensor includes a first substrate 579 with
at least one working electrode 580, as illustrated in FIG. 22A. The
sensor also includes a spacer 581 with a channel 582 formed in the
spacer 581, as shown in FIG. 221. The sensor further includes a
second substrate 583 with two counter electrodes 584, 585, as shown
in FIG. 22C (inverted with respect to FIGS. 22A and 22B). A venting
aperture 586 is cut typically through all of the layers and extends
from a side of the sensor. In some embodiments, the venting
aperture and the front portion 587 of the sensor are simultaneously
cut with a reproducible distance between the venting aperture and
the front portion 587 of the sensor to provide a reproducible
length for the channel 582 and the working electrode 580. FIGS.
22A, 22B, and 22C also illustrate another feature that can be used
with any sensor arrangement. An indentation 588 may be formed at
the filling opening of the channel 582 to facilitate the drawing of
fluid into the sensor. In this configuration, the fluid is not
provided with a flat face, but rather an indented face that may aid
in wicking or capillary filling of the channel (i.e., sample
chamber). This configuration may also reduce the likelihood that
the user of the sensor will block the channel during collection of
the sample. A flat faced sensor might be blocked by pressing the
tip of the sensor edgewise against the skin.
FIGS. 23A, 23B, and 23C illustrate another example of a
side-filling sensor arrangement. FIG. 23A illustrates a first
substrate 640 with a working electrode 642. FIG. 23B illustrates a
spacer 644 defining a channel 646. FIG. 23C (inverted with respect
to FIGS. 23A and 23B) illustrates a second substrate 648 with three
counter (or counter/reference) electrodes 650, 652, 654. This
sensor can be formed by making straight cuts of the substrates. The
sensors can be made adjacent to one another, as illustrated in FIG.
31A, which may produce less waste material. The length of the
channel 646 is typically defined by the two parallel cuts along the
sides 656, 658 of the sensors. Another optional processing
advantage, particularly if the sensor are formed adjacent to each
other, is that the redox mediator and/or second electron transfer
agent can be disposed in the channel by striping a continuous
stream of these components along a row or column of adjacent
sensors. This may result in better efficiency and less waste of the
redox mediator and/or second electron transfer agent, as compared
to other techniques, such as individually placing these components
within the individual channels.
FIGS. 24A, 24B, and 24C illustrate another sensor configuration.
This sensor includes a first substrate 600 with at least one
working electrode 602, as illustrated in FIG. 24A. The sensor also
includes a spacer 604 with a channel 606 formed in the spacer 604,
as shown in FIG. 24B. The sensor further includes a second
substrate 608 with two counter electrodes 610, 612, as shown in
FIG. 24C (inverted with respect to FIGS. 24A and 24B). The sensor
may also include, for example, an indicator, such as a slot 614 or
an extension 616 from the body of the sensor that indicates to the
user which side should be placed adjacent to the sample. This may
be particularly important where the sensor reading is only correct
when sample enters from a particular side.
FIG. 24B also illustrates another optional feature that may be used
in any of the sensor configurations. In this illustration, the
sample chamber 606 is not formed using straight lines, but there is
an expanded region 618 within the sample chamber. This permits
larger sample chambers without forming larger openings. This
expanded region can be formed as any shape including circular,
square, rectangular, and other regular and irregular shapes.
FIG. 25 is an example of an assembled sensor illustrating another
alternative sensor arrangement for a side-fill sensor 620. This
sensor includes extensions 622 from the sensor body 624 to indicate
to a user where the openings for the sample chamber 626 are
provided.
One optional feature is illustrated in FIG. 32 which is an edge-on
view of the sensor from the inside of the meter. FIG. 32
illustrates a first substrate 1120 and a second substrate 1130 that
extend into the meter from the remainder of the sensor 1100 (i.e.,
portion 1140 is recessed with respect to substrates 1120 and 1130
in FIG. 32). Examples of this configuration are illustrated in
FIGS. 18A-18C and 24A-24C. Typically, the sensor 1100 is coupled to
a meter 1110 that includes contact pads (not shown) that contact
the contact regions (e.g., regions 503, 511, and 513 in FIGS. 18A
and 18C) of the electrodes of the sensor 1100. The end of the
sensor 1100 which contains the contact regions can be slid into the
meter 1110. It is typically important that the contact pads of the
meter 1110 make contact with the correct contact regions of the
sensor so that the working electrode and counter electrode(s) are
correctly coupled to the meter. In some instances, the sensor is
configured so that the contact region for the working electrode on
the first substrate 1120 has a different width, w1, than width, w2,
for the contact region of the second substrate 1130 carrying the
counter electrode(s). Examples of electrode configurations with
this structure are provided in FIGS. 18A-18C and 24A-24C. To ensure
proper insertion of the sensor 1100 into the meter 1110, the meter
1110 may include a raised area 1140 that prevents or hinders the
insertion of the sensor in an improper direction. For example, the
width, w2, of the contact region of the second substrate 1130 may
be wider than the width, w1, of the contact region of the first
substrate 1120, as illustrated in FIG. 32. In this instance, the
raised area 1140 is positioned to allow sensor 1100 to be slid into
the meter so that the first substrate 1120 is next to the surface
1150 from which the raised area 1140 protrudes, but would prevent
or hinder having the second substrate 1130 next to the surface 1150
from which the raised area 1140 protrudes. Objects other than a
raised area can also be used to guide the user in correct
introduction of the sensor into the meter.
Integrated Sample Acquisition and Analyte Measurement Device
Many approaches are known in the art for acquiring and/or
transporting a small sample from the body to a sensor. These
include, for example, U.S. Pat. Nos. 5,746,217; 5,820,570;
5,857,983; and 5,879,311, incorporated herein by reference. Any of
these sample acquisition and/or transporting methods may be
employed with the sensor of the current invention.
In a preferred embodiment of the invention, an analyte measurement
device 52 constructed according to the principles of the present
invention includes a sensor 20, as described hereinabove, combined
with a sample acquisition apparatus 50 to provide an integrated
sampling and measurement device. The sample acquisition apparatus
50 illustrated in FIG. 6, includes, for example, a skin piercing
member 54, such as a lancet, attached to a resilient deflectable
strip 56 (or other similar device, such as a spring) which may be
pushed to inject the lancet into a patient's skin to cause blood
flow.
The resilient strip 56 is then released and the skin piercing
member 54 retracts. Blood flowing from the area of skin pierced by
member 54 can then be transported, for example, by the wicking
action of sorbent material 34, into sensor 20 for analysis of the
analyte. The analyte measurement device 52 may then be placed in a
reader, not shown, which connects a coulometer or other
electrochemical analysis equipment to the electrode tabs 23, 25 to
determine the concentration of the analyte by electroanalytical
means. Preferably, the analyte measurement device is enclosed
within the reader when connected to the coulometer or other
electrochemical analysis equipment.
In a preferred embodiment, the integrated sample acquisition and
analyte measurement device comprises a lancing instrument that
holds a lancet and measurement strip. The lancing instrument
preferably requires active cocking. By requiring the user to cock
the device prior to use, the risk of inadvertently triggering the
lancet is minimized.
Preferably, the lancing instrument is automatically triggered when
the lancing instrument is pressed firmly against the skin with an
adequate amount of pressure. As is already known in the art, a
larger sample of body fluid such as blood or interstitial fluid is
expressed when pressure is applied around a site where a hole has
been created the skin. For example, see the above-mentioned U.S.
patents to Integ and Amira as well as the tip design of the lancing
instruments sold by Becton Dickenson. All of these lancing devices
have a protruding ring that surrounds the lancing site to create
pressure that forces sample out of the wound. However, all of these
devices require the user to apply adequate pressure to the wound
site to express the sample, and all of the lancing instruments are
triggered by a button push by the user. Design of an appropriate
pressure trigger is well-known to one skilled in the art.
Preferably, the lancing instrument will also permit the user to
adjust the depth of penetration of the lancet into the skin. Such
devices are already commercially available from companies such as
Bochringer Mannheim and Palco. This feature allows users to adjust
the lancing device for differences in skin thickness, skin
durability, and pain sensitivity across different sites on the body
and across different users.
In a more preferred embodiment, the lancing instrument and the test
reader are integrated into a single device. To operate the device
the user need only insert a disposable cartridge containing a
measurement strip and lancing device into the integrated device,
cock the lancing instrument, press it against the skin to activate
it, and read the result of the measurement. Such an integrated
lancing instrument and test reader simplifies the testing procedure
for the user and minimizes the handling of body fluids.
FIG. 26 illustrates another example of an integrated sample
acquisition and sensor device 700. The integrated sample
acquisition and sensor device 700 includes a housing 702, a skin
piercing member (e.g., a lancet) 704, a piercing/collecting
aperture 706, an optionally removable sensor 708, a sensor guide
710, and a retraction mechanism 714 for the skin piercing member.
This device 700 can be designed for reuse (e.g., by making the skin
piercing member 704 and sensor 708 removable) or for single
use.
The housing 702 may be formed of a variety of materials including
metal and plastic. The housing 702 may include a hinge 716 or other
configuration (e.g., adhesive or interlocking parts) for holding
portions of the housing together.
The piercing/collecting aperture 706 is provided in the housing 702
to allow the skin piercing member 704 to extend through the
aperture 706 and pierce the skin of a user, thereby causing blood
(or other body fluid) flow. The sensor 708 also extends to the edge
or out of the aperture 706 to collect the blood (or other body
fluid) through an opening (not shown) in the tip of the sensor.
This may allow the user to pierce the skin and collect the fluid
sample without moving the device 700. Alternatively, separate
apertures may be provided for the skin piercing member 704 and
sensor 708. The sensor guide may be formed in the housing 702 or
added to the housing to guide the sensor 708 into place if the
sensor is inserted into and through the housing and/or to support
the sensor within the housing and during sample collection.
The skin piercing member 704 may include an actuator (not shown)
that includes a mechanism that allows for cocking and releasing the
skin piercing member 704 or the skin piercing member may be
actuated externally. For example, a sensor reader (not shown) or
other device may be coupled to the sample acquisition and sensor
device, the sensor reader or other device including a mechanism
that cocks and/or releases the skin piercing member 704.
The retraction mechanism 714 of the device 700 may be, for example,
a spring or resilient metal strip that retracts the skin piercing
member 704 back into the housing after piercing the skin of the
user. This may allow for unobstructed collection of the sample
and/or prevent further piercing of the skin of the user or others
to reduce or prevent contamination or infection caused by transfer
of body fluids or other harmful agents. Alternatively, retraction
of the skin piercing member may be accomplished using an external
device or apparatus.
One example of operation includes cocking the skin piercing member
704 and then releasing the skin piercing member 704 so that it
extends out of the housing 702 through the piercing/collecting
aperture 706 and pierces the skin of the user. The skin piercing
element 704 optionally pushes the sensor out of the way while
extending out of the housing. The skin piercing element 704 is
retracted back within the housing 702 using the retraction
mechanism 714. Upon retraction of the skin piercing element, the
sensor collects a sample fluid from the pierced skin through an
opening in the sensor 708.
If a sensor reader is used, the sensor reader may also be
configured to couple with a contact end of the sensor. The sensor
reader may include a potentiostat or other component to provide a
potential and/or current for the electrodes of the sensor. The
sensor reader may also include a processor (e.g., a microprocessor
or hardware) for determining analyte concentration from the sensor
signals. The sensor reader may include a display or a port for
coupling a display to the sensor. The display may display the
sensor signals and/or results determined from the sensor signals
including, for example, analyte concentration, rate of change of
analyte concentration, and/or the exceeding of a threshold analyte
concentration (indicating, for example, hypo- or hyperglycemia).
This sensor reader may be used in conjunction with the integrated
sample acquisition and sensor device or the sensor reader may be
used with the sensor alone, the contacts of the sensor making
connection with contacts in the sensor reader.
Operation of the Sensor
An electrochemical sensor of the invention may be operated with or
without applying a potential. In one embodiment, the
electrochemical reaction occurs spontaneously and a potential need
not be applied between the working and counter electrodes.
In another embodiment, a potential is applied between the working
and counter electrodes. Yet the potential does not need to remain
constant. The magnitude of the required potential is dependent on
the redox mediator. The potential at which the electrode poises
itself, or where it is poised by applying an external bias, and
where the analyte is electrolyzed is typically such that the
electrochemical reaction is driven to or near completion, but it
is, preferably, not oxidizing enough to result in significant
electrochemical reaction of interferents, such as urate, ascorbate,
and acetaminophen, that may affect the signal measured. For
non-leachable redox mediators, the potential is typically between
about -350 mV and about +400 mV versus the standard calomel
electrode (SCE). Preferably, the potential of the redox mediator is
more negative than +100 mV, more preferably the potential is more
negative than 0 mV, and most preferably the potential is about -150
mV versus SCE.
When an external potential is applied, it may be applied either
before or after the sample has been placed in the sample chamber.
If the measurement zone comprises only a portion of the sample
chamber then the potential is preferably applied after the sample
has come to rest in the sample chamber to prevent electrolysis of
sample passing through the measurement zone as the sample chamber
is filling. Alternatively, in the case where the measurement zone
comprises most or all of the sample chamber, the potential,
optionally, may be applied before or during the filling of the
sample chamber without affecting the accuracy of the assay. When
the potential is applied and the sample is in the measurement zone,
an electrical current will flow between the working electrode and
the counter electrode. The current is a result, at least in part,
of the electrolysis of the analyte in the sample. This
electrochemical reaction occurs via the redox mediator and the
optional second electron transfer agent. For many biomolecules, B,
the process is described by the following reaction equations:
##STR00009##
Biochemical B is oxidized to C by redox mediator species A in the
presence of an appropriate enzyme. Then the redox mediator A is
oxidized at the electrode. Electrons are collected by the electrode
and the resulting current is measured. The measured current may
also include a background current resulting in a measured
background charge, due, at least in part, to the shuttling of a
diffusible redox mediator between the working electrode and the
counter electrode. This background current can be minimized or
accounted for, as described above.
As an example, one sensor of the present invention is based on the
reaction of a glucose molecule with two
[Os(dmo-phen).sub.2(NMI)Cl].sup.2+ cations, where dmo-phen is
4,8-dimethoxy phenanthroline and NMI is N-methyl-imidazole, in the
presence of glucose oxidase to produce two
[Os(dmo-phen).sub.2(NMI)Cl].sup.+ cations, two protons, and an
oxidation product of glucose, for example, gluconolactone or
another ketone. The amount of glucose present is assayed by
electrooxidizing the [Os(dmo-phen).sub.2(NMI)Cl].sup.+ cations to
[Os(dmo-phen).sub.2(NMI)Cl].sup.2+ cations and measuring the total
charge passed.
Those skilled in the art will recognize that there are many
different reactions that will provide the same result; namely the
electrolysis of an analyte through a reaction pathway incorporating
a redox mediator. Equations (1) and (2) are a non-limiting example
of such a reaction.
Coulometry
In a preferred embodiment of the invention, coulometry is used to
determine the concentration of the analyte. This measurement
technique utilizes current measurements obtained at intervals over
the course of the assay, to determine analyte concentration. These
current measurements are integrated over time to obtain the amount
of charge, Q, passed to or from the electrode. Q is then used to
calculate the concentration of the analyte (C.sub.A) by the
following equation (when the redox mediator is non-leachable):
C.sub.A=Q/nFV (3a) where n is the number of electron equivalents
required to electrolyze the analyte, F is Faraday's constant
(approximately 96,500 couiombs per equivalent), and V is the volume
of sample in the measurement zone. When using a diffusible
mediator, the concentration of the analyte can be obtained from the
following equation: C.sub.A+(Q.sub.tot-Q.sub.back)/nFV (3b) where
Q.sub.tot is the total charge transferred during the measurement
and Q.sub.back is the amount of charge transferred that was not due
to the analyte, e.g., charge transferred by the shuttling of the
diffusible mediator between the working electrode and the counter
electrode. In at least some instances, the sensor is constructed so
that the background charge is at most 5 times the size of the
charge generated by electrolysis of an amount of analyte.
Preferably, the background signal is at most 200%, 100%, 50%, 25%,
10%, or 5% of the charge generated by electrolysis of the
analyte.
One example of a method for determining the ratio of background
signal to signal generated by electrolysis of the analyte is
described as follows for the facing electrode pairs. If the
shuttling of the redox mediator is not disabled by the applied
potential, the charge that results from the shuttling of the redox
mediator may be represented by the following formula:
Q.sub.back=(AFD.sub.MC.sub.M/d)(tn.sub.M) where A is the area of
the working electrode; F is Faraday's constant (96,500
coulomnbs/equivalent); D.sub.M is the effective diffusion
coefficient of the redox mediator, C.sub.M is the concentration of
the redox mediator in the measurement zone; d is the distance
separating facing electrodes; t is the amount of time for the
measurement; and n.sub.M is the number of electrons gained or lost
by the redox mediator.
Additionally, the charge of the analyte, for example, glucose, when
the analyte is electrooxidized to about 90% completion in the
measurement period may be represented by the following formula:
Q.sub.G=Ad(0.90)C.sub.Gn.sub.GF where A is the area of the working
electrode; d is the distance separating facing electrodes; C.sub.G
is the concentration of glucose; n is the number of electrons
needed to electrolyze the analyte (e.g., 2 electrons per glucose
molecule); and F is Faraday's constant. When C.sub.G is 5 mM (or
5.times.10.sup.-6 moles/cm.sup.3), t is 60 seconds, n.sub.G is 2,
and n.sub.M is 1, the ratio of charge from the redox mediator to
the charge from electrooxidation of the analyte may be represented
by the following formula:
Q.sub.Back/Q.sub.G=(D.sub.MC.sub.M/d.sup.2)(tn.sub.M/(0.9n.sub.G-
C.sub.G))=(D.sub.MC.sub.M/d.sup.2).times.(6.7.times.10.sup.6) For
example, if the ratio of Q.sub.Back/Q.sub.G is 5, then (D.sub.M
C.sub.M)/d.sup.2 is 7.5.times.10.sup.-7 moles/(cm.sup.3 sec). Also
for example, if the ratio of Q.sub.Back/Q.sub.G is 1, then (D.sub.M
C.sub.M)/d.sup.2 is 1.5.times.10.sup.-7 moles/(cm.sup.3 see). Still
another example, if the ratio is 0.1, then (D.sub.M
C.sub.M)/d.sup.2 is 1.5.times.10.sup.-8 moles/(cm.sup.3 see). Thus,
depending on the ratio desired, a sensor may be configured to have
the desired ratio by choosing D.sub.M, C.sub.M, and d accordingly.
For example, the concentration of the redox mediator may be reduced
(i.e., C.sub.M may be reduced). Alternatively, or additionally, the
diffusion of the redox mediator may be reduced by, for example,
having a barrier to the flow of the diffusible mediator to the
counter electrode (i.e., reduce the effective diffusion coefficient
of the redox mediator--D.sub.M). Other sensor configurations are
also suitable for controlling the ratio of background signal to
signal generated by the analyte and will be described below.
The background charge, Q.sub.back, can be accounted for in a
variety of ways. Q.sub.Back can be made small, for example, by
using only limited amounts of diffusible redox mediator; by
providing a membrane over the counter electrode that limits
diffusion of the redox mediator to the counter electrode; or by
having a relatively small potential difference between the working
electrode and the counter electrode. Other examples of sensor
configurations and methods suitable for reducing Q.sub.back include
those already described such as sensors having a redox mediator
reaction rate at the working electrode that is significantly faster
than that at the counter electrode; immobilizing the redox mediator
on the working electrode; having the redox mediator become
immobilized on the counter or counter/reference electrode upon its
reaction at the counter or counter/reference electrode; or slowing
the diffusion of the redox mediator.
Alternatively, the sensor may be calibrated individually or by
batch to determine a calibration curve or a value for Q.sub.back.
Another option is to include a second electrode pair that is
missing an item necessary for electrolysis of the analyte, such as,
for example, the second electron transfer agent, so that the entire
signal from this second electrode pair corresponds to
Q.sub.back.
For coulometric measurements, at least 20% of the analyte is
electrolyzed. Preferably at least 50%, more preferably at least
80%, and even more preferably at least 90% of the analyte is
electrolyzed. In one embodiment of the invention, the analyte is
completely or nearly completely electrolyzed. The charge can then
be calculated from current measurements made during the
electrochemical reaction, and the concentration of the analyte is
determined using equation (3a) or (3b). The completion of the
electrochemical reaction is typically signaled when the current
reaches a steady-state value. This indicates that all or nearly all
of the analyte has been electrolyzed. For this type of measurement,
at least 90% of the analyte is typically electrolyzed, preferably,
at least 95% of the analyte is electrolyzed and, more preferably,
at least 99% of the analyte is electrolyzed.
For coulometry, it is typically desirable that the analyte be
electrolyzed quickly. The speed of the electrochemical reaction
depends on several factors, including the potential that is applied
between the electrodes and the kinetics of reactions (1) and (2).
(Other significant factors include the size of the measurement zone
and the presence of sorbent in the measurement zone.) In general,
the larger the potential, the larger the current through the cell
(up to a transport limited maximum) and therefore, the faster the
reaction will typically occur. However, if the potential is too
large, other electrochemical reactions may introduce significant
error in the measurement. Typically, the potential between the
electrodes as well as the specific redox mediator and optional
second electron transfer agent are chosen so that the analyte will
be almost completely electrolyzed in less than 5 minutes, based on
the expected concentration of the analyte in the sample.
Preferably, the analyte will be almost completely electrolyzed
within about 2 minutes and, more preferably, within about 1
minute.
In another embodiment of the invention, the analyte is only
partially electrolyzed. The current is measured during the partial
reaction and then extrapolated using mathematical techniques known
to those skilled in the art to determine the current curve for the
complete or nearly complete electrolysis of the analyte.
Integration of this curve yields the amount of charge that would be
passed if the analyte were completely or nearly completely
electrolyzed and, using equation (3a) or (3b), the concentration of
the analyte is calculated.
Although coulometry has the disadvantage of requiring the volume of
the measured sample be known, coulometry is a preferred technique
for the analysis of the small sample because it has the advantages
of, for example, no temperature dependence for the measurement, no
enzyme activity dependence for the measurement, no redox-mediator
activity dependence for the measurement, and no error in the
measurement from depletion of analyte in the sample. As already
described above, coulometry is a method for determining the amount
of charge passed or projected to pass during complete or nearly
complete electrolysis of the analyte. One coulometric technique
involves electrolyzing the analyte on a working electrode and
measuring the resulting current between the working electrode and a
counter electrode at two or more times during the electrolysis. The
electrolysis is complete when the current reaches a steady state.
The charge used to electrolyze the sample is then calculated by
integrating the measured currents over time and accounting for any
background signal. Because the charge is directly related to the
amount of analyte in the sample there is no temperature dependence
of the measurement. In addition, the activity of the enzyme does
not affect the value of the measurement, but only the time required
to obtain the measurement (i.e., less active enzyme requires a
longer time to achieve complete electrolysis of the sample) so that
decay of the enzyme over time will not render the analyte
concentration determination inaccurate. And finally, the depletion
of the analyte in the sample by electrolysis is not a source of
error, but rather the objective of the technique. (However, the
analyte need not be completely electrolyzed if the electrolysis
curve is extrapolated from the partial electrolysis curve based on
well-known electrochemical principles.)
Non-Coulometric Assays
Although coulometric assays are useful, those skilled in the art
will recognize that a sensor of the invention may also utilize
potentiometric, amperometric, voltammetric, and other
electrochemical techniques to determine the concentration of an
analyte in a sample. The measurements obtained by these
non-coulometric methods may not be temperature independent as are
coulometric measurements.
In addition, the measurements obtained by these non-coulometric
electrochemical techniques may be sensitive to the amount of active
enzyme provided in the sensor. If the enzyme deactivates or decays
over time, the resulting measurements may be affected. This may
limit the shelf life of such sensors unless the enzyme is very
stable.
Finally, the measurements obtained by non-coulometric
electrochemical techniques, such as steady-state amperometry, may
be negatively affected if a substantial portion of the analyte
and/or redox mediator is electrolyzed during the measurement
period. An accurate steady-state measurement may not be obtainable
unless there is sufficient analyte and/or redox mediator so that
only a relatively small portion of the analyte and/or redox
mediator is electrolyzed during the measurement process. This may
be challenging in a sample size of no more than 1 .mu.l.
It may be desirable in some instances to utilize non-coulometric
assays, such as amperometric or potentiometric measurement
techniques. For example, coulometry requires that the volume of the
measured sample be known. And, the volume of the sample in the
measurement zone of a small volume sensor (i.e., no more than one
microliter) may be difficult to accurately reproduce if the
manufacturing tolerances of one or more dimensions of the
measurement zone have significant variances.
As described for coulometric measurements, the background signal
resulting from the shuttling of the redox mediator between the
electrodes can be a source of measurement error in amperometric or
potentiometric assays of samples of no more than 1 .mu.L in thin
layer electrochemical cells. In general, it is desirable that the
mediator does not shuttle between a pair of electrodes more than
ten times in the period of the measurement, preferably not more
than once, and more preferably not more than 0.1 times, on average.
To decrease error arising from background signal, methods and
sensor configurations similar to, and in some cases identical to,
those used for coulometric measurements may be used. Examples
include all of the methods and structures described above, such as
performing the electrochemical assay at relatively low applied
potential, electrooxidizing the analyte at negative applied
potentials or electroreducing the analyte at positive applied
potentials, using a counter electrode at which the redox mediator
reacts relatively slowly (particularly as compared to the reaction
of the redox mediator at the working electrode), and/or using a
redox mediator that undergoes an irreversible reaction at the
counter electrode. Other examples are discussed below.
As described for coulometric measurements, it is preferred that the
sensor be designed and operated so that the background signal is at
most five times the size of the signal generated by electrolysis of
the analyte. Preferably, the background signal is at most 200%,
100%, 50%, 25%, 10%, or 5% of the signal generated by electrolysis
of an amount of analyte. The amount of analyte against which the
background signal is compared is described above in the section
entitled "Background Signal." In the case of amperometry, the
signal generated by electrolysis of an amount of analyte is the
current at the time or times at which the measurement is taken. In
the case of potentiometry, the signal generated by electrolysis of
an amount of analyte is the potential at the time or times at which
the measurement is taken.
Under a given set of operating conditions, for example,
temperature, cell geometry, and electrode size, the magnitude of
the background current, I.sub.back, is given by the following
expression: i.sub.back=KC.sub.MD.sub.M/d where: K is a
proportionality constant; C.sub.M is the concentration of the
mediator in the measurement zone; D.sub.m is the effective
diffusion coefficient of the mediator in the measurement zone under
normal operating conditions; and d is the distance between the
electrodes.
It is desirable to reduce background current for non-coulometric
assays. The sensor configurations and methods described above are
generally useful and include, for example, using low concentrations
of the redox mediator and/or the second electron transfer agent
(e.g., enzyme) relative to the concentration of the analyte and/or
using a large redox mediator having a relatively low effective
diffusion constant. Other useful methods described above include
methods for reducing the diffusion of the redox mediator by, for
example, having a barrier (e.g., a charged or polar barrier) to the
flow of the diffusible mediator or using a redox mediator having a
relatively low effective diffusion constant.
In some instances, the effective diffusion coefficient is no more
than about 1.times.10.sup.-6 cm.sup.2/sec, no more than about
1.times.10.sup.-7 cm.sup.2/sec, or no more than about
1.times.10.sup.-8 cm.sup.2/sec. Moreover, in some cases, the
product of C.sub.MD.sub.M (the concentration of redox mediator
times the effective diffusion coefficient) is no more than about
1.times.10.sup.-12 moles/cmsec, no more than about
1.times.10.sup.-13 moles/cmsec, or no more than about
1.times.10.sup.-14 moles/cmsec.
The following provides a specific example for the case of an
amperometric measurement of glucose carried out for 60 seconds
during which time 10% of the glucose is electrolyzed in a 1
microliter cell with facing electrodes separated by a distance of
d=0.01 cm. If the measurement was carried out under the following
conditions: a glucose concentration of, C.sub.G=5 mM (or
5.times.10.sup.-6 moles/cm.sup.3), an area of A=0.1 cm.sup.2, a
number of electrons from the redox mediator of n.sub.M=1, and a
number of electrons from glucose n.sub.G=2, then the background
current generated by the redox mediator and by the glucose is
determined as follows.
.times..times..times..times..times..times..times..times..times..times.
##EQU00001##
.times..function..times..times..times..times..times..times..times..times.-
.times..times. ##EQU00001.2##
Thus if i.sub.back/i.sub.G=5, the value of C.sub.MD.sub.M equal
8.34.times.10.sup.-12 moles/cm.sup.2 sec. As another example, if
i.sub.back/i.sub.G=0.5, the value of C.sub.MD.sub.M equal
8.34.times.10.sup.-13 moles/cm.sup.2 sec. Additionally if
i.sub.back/i.sub.G=0.05, the value of C.sub.MD.sub.M equal
8.34.times.10.sup.-14 moles/cm.sup.2 sec.
In some amperometric or potentiometric embodiments, the redox
mediator circulation is decreased by separating the working
electrode from the counter or counter/reference electrode such that
the distance through which the redox mediator would diffuse during
the measurement period is no greater than, for example, the
distance between the electrodes. A redox mediator can diffuse a
distance equal to (D.sub.mt).sup.1/2, where D.sub.m is the
effective diffusion coefficient for the medium between the
electrodes and t is time. For a measurement time period of 30
seconds and a redox mediator with effective diffusion coefficient
between 10.sup.-5 and 10.sup.-6 cm.sup.2/second, the electrodes
should be separated by at least 100 .mu.m, preferably at least 200
.mu.m, and even more preferably at least 400 .mu.m.
One method of separating the working and counter electrodes is to
use a thicker spacer between the electrodes. One alternative method
is illustrated in FIG. 27. In this embodiment, the working
electrode 740 is disposed on a first substrate 742 and the counter
electrode 744 is disposed on a second substrate 746 (alternatively,
the electrodes may be disposed on the same substrate). The working
electrode 742 and the counter electrode 744 are offset so that the
effective distance, d, between the two electrodes is greater than
the thickness, w, of the spacer layer 748. In one embodiment, the
distance between the electrodes, d, is selected to be in the range
of 25 to 1000 .mu.m, 50 to 500 .mu.m, or 100 to 250 .mu.m.
Additionally or alternatively, in the case of steady-state
amperometry and potentiometry, background signal may be controlled
by limiting the rate of electrolysis such that the rate is slow
enough to prevent the analyte concentration from decreasing by more
than about 20%, 10%, or 5% or less, during a measurement period,
e.g., 30 second, 1 minute, 5 minutes, or 10 minutes. In some
instances, to control the rate of electrolysis the concentration or
activity of the second electron transfer agent may be reduced
and/or the working electrode area may be reduced.
For example, the second electron transfer agent can be an enzyme
and the enzyme activity can be a limiting factor for the
electrolysis rate. If, for example, the analyte concentration is 5
mM glucose (i.e., 5.times.10.sup.-9 moles of glucose in 1 .mu.l)
and no more than 10% of the glucose (5.times.10.sup.-10 moles) is
to be electrooxidized during a 30-second measurement period, the
current should not exceed 3.3.times.10.sup.-6 amperes for 1 .mu.L.
One unit of an enzyme is that amount of the enzyme which catalyzes
electrolysis of 1 .mu.mole of its substrate in 60 seconds at pH of
7.4 at 37.degree. C. in HEPES buffer. Accordingly, for glucose, a
current of up to 3.3.times.10.sup.-3 amperes in 1 cm.sup.3 (i.e., 1
mL) can be generated. Therefore, the maximum amount of enzyme used
in a sensor that limits the amount of electrolysis by controlling
the amount of enzyme should be 1 unit/cm.sup.3 or less.
The rate of electrolysis may also be limited by using a relatively
small working electrode area. When the working electrode area is
sufficiently small (e.g., no more than about 0.01 cm.sup.2, no more
than about 0.0025 cm.sup.2, or no more than about 0.001 cm.sup.2),
then radial diffusion of analyte to the electrode may result in a
steady-state current, at a constant applied potential, that is
representative of the analyte concentration. For circular
electrodes, the appropriate surface area may be achieved using an
electrode with a radius of no more than 60 .mu.m, no more than 30
.mu.m, or no more than 20 .mu.m. Radial diffusion of the analyte
includes transport of analyte from all directions and not just from
the direction normal to the electrode surface and can, therefore,
reduce or prevent depletion of analyte near the electrode surface.
A small electrode on a planar surface permits radial diffusion. In
a sensor having a larger surface area electrodes, the transport of
analyte to the electrode may be modeled as semi-infinite linear
diffusion instead of radial diffusion. Thus, the transport of the
analyte to the electrode is dominated by diffusion from the
direction normal to the electrode surface. As a result, the reduced
transport rate is typically unable to overcome the depletion of
analyte near the electrode surface, and at constant applied
potential the current decreases with time, t, according to
t.sup.-1/2.
For a potentiometric assay of the type proposed by Yarnitzky and
Heller, J. Phy Chem., 102:10057-61 (1998), in which the potential
varies linearly with the analyte concentration, the concentration
of the analyte and/or redox mediator in a particular oxidation
state should vary no more than about 20% during the assay. If the
concentration varies by more than 20%, then the diffusion of the
analyte or redox mediator should be controlled by, for example,
controlling temperature and/or volume of the sample chamber and/or
measurement zone.
While this description has described electrolysis of an analyte,
one skilled in the art would recognize that the same devices and
techniques would also be suitable for measurements of the average
oxidation state of the mediator, such as, for example, in Cottrell
types of reactions.
Air-oxidizable Redox Mediators
In a sensor having a redox mediator, a potential source of
measurement error is the presence of redox mediator in an unknown
mixed oxidation state (i.e., mediator not reproducibly in a known
oxidation state). The charge passed when the redox mediator is
electrooxidized or electroreduced at the working electrode is
affected by its initial oxidation state. Referring to equations (1)
and (2) discussed above under the section entitled "Operation of
the Sensor," the current not attributable to the oxidation of
biochemical B will flow because of electrooxidation of that portion
of the redox mediator, A, that is in its reduced form prior to the
addition of the sample. Thus, it may be important to know the
oxidation state of the analyte prior to introduction of the sample
into the sensor. Furthermore, it is desirable that all or nearly
all of the redox mediator have the same state or extent of
oxidation prior to the introduction of the sample into the
sensor.
Each redox mediator has a reduced form or state and an oxidized
form or state. It is preferred that the amount of redox mediator in
the reduced form prior to the introduction of sample be
significantly smaller than the expected amount of analyte in a
sample in order to avoid a significant background contribution to
the measured current. In this embodiment of the invention, the
molar amount of redox mediator in the reduced form prior to the
introduction of the analyte is preferably no more than, on a
stoichiometric basis, about 10%, and more preferably no more than
about 5%, and most preferably no more than 1%, of the molar amount
of analyte for expected analyte concentrations. (The relative molar
amounts of analyte and redox mediator are compared based on the
stoichiometry of the applicable redox reaction. If, for example,
two moles of redox mediator are needed to electrolyze one mole of
analyte, then the molar amount of redox mediator in the reduced
form prior to introduction of the analyte is preferably no more
than 20% and more preferably no more than about 10% and most
preferably no more than about 2% of the molar amount of analyte for
expected analyte concentrations.) Methods for controlling the
amount of reduced mediator are discussed below.
In another aspect of the invention, it is preferred that the ratio
of the amounts of oxidized redox mediator to reduced redox
mediator, prior to introduction of the sample in the sensor, be
relatively constant between similarly constructed sensors. Any
deviation from holding the ratio relatively constant may increase
the scatter of the results obtained for the same sample with
multiple similarly made sensors. For this aspect of the invention,
the percentage of the redox mediator in the reduced form prior to
introduction of the sample in the sensor varies by no more than
about 20% and preferably no more than about 10% between similarly
constructed sensors.
One method of controlling the amount of reduced redox mediator
prior to the introduction of the sample in the sensor is to provide
an oxidizer to oxidize the reduced form of the mediator. One of the
most convenient oxidizers is O.sub.2. Oxygen is usually readily
available to perform this oxidizing function. Oxygen can be
supplied by exposing the sensor to air. In addition, most polymers
and fluids absorb O.sub.2 from the air unless special precautions
are taken. Typically, at least 90% of an air-oxidizable (i.e.,
O.sub.2 oxidizable) mediator in the solid state is in the oxidized
state upon storage or exposure to air for a useful period of time,
e.g., one month or less, and preferably, one week or less, and,
more preferably, one day or less. The air oxidation may take place
in either the solid state or as a solution stored for a time period
sufficient to air oxidize the mediator before deposition on the
sensor. In the case of air oxidizable redox mediators in solution,
it is desirable that the time required to achieve at least 80%,
preferably at least 90%, oxidation of the redox mediator is at
least 10 times the expected duration of the assay and is also less
than the pot life of the solution. Preferably, at least 80%, more
preferably at least 90%, of the redox mediator is air oxidized in
less than 1 week, preferably, in less than 1 day, more preferably,
in less than 8 hours, and even more preferably, in less than 1
hour.
While it is desirable to bring the mediators of the sensors
manufactured in a single batch to the same state or extent of
oxidation, it is not necessary that the mediator be completely
oxidized to the higher-valent state. Additionally, it is desirable
that the air oxidation of the dissolved redox mediator should not
be so fast that air-oxidation during the assay can interfere with
or introduce error into the measurements.
Suitable mediators which are both air-oxidizable (i.e.,
O.sub.2-oxidizable) and have electron transfer capabilities have
been described hereinabove. One particular family of useful
mediators are osmium complexes which are bound to electron-rich
nitrogen-containing heterocycles or a combination of electron-rich
nitrogen-containing heterocycles and halides. Electron-rich
nitrogen-containing heterocycles include, but are not limited to,
imidazole derivatives and pyridine or phenanthroline derivatives
that contain electron-donating substituents such as alkyl, alkoxy,
amino, alkylamino, amido and mercapto groups. Preferably, the
osmium complexes have no more than one halide coordinated to the
metal, so that the mediators are overall positively charged and
thus are water soluble. An example is osmium complexed with mono-,
di-, and polyalkoxy-2,2'-bipyridine. Other examples include mono-,
di-, and polyalkoxy-1,10-phenanthroline, where the alkoxy groups
have a carbon to oxygen ratio sufficient to retain solubility in
water, are air-oxidizable. These osmium complexes typically have
two substituted bipyridine or substituted phenanthroline ligands,
the two ligands not necessarily being identical. These osmium
complexes are further complexed with a monomeric or polymeric
ligand with one or more nitrogen-containing heterocycles, such as
pyridine and imidazole. Preferred polymeric ligands include
poly(4-vinyl pyridine) and, more preferably, poly(1-vinyl
imidazole) or copolymers thereof.
[Os[4,4'-dimethoxy-2,2'-bipyridine].sub.2Cl].sup.+/+2 complexed
with a poly(1-vinyl imidazole) or poly(4-vinyl pyridine) has been
shown to be particularly useful as the Os.sup.+2 cation is
oxidizable by O.sub.2 to Os.sup.+3. Similar results are expected
for complexes of
[Os(4,7-dimethoxy-1,10-phenanthroline).sub.2Cl].sup.+/+2, and other
mono-, di-, and polyalkoxy bipyridines and phenanthrolines, with
the same polymers. Other halogen groups such as bromine may be
substituted for chlorine. Similar results are also expected for
complexes comprising the following structures, as specified
above:
##STR00010##
A complication associated with air-oxidizable mediators arises if
the air oxidation of the redox mediator is so fast that a
substantial portion of the analyte-reduced redox mediator is
oxidized by O.sub.2 during an analyte assay. This will result in an
inaccurate assay as the amount of analyte will be underestimated
because the mediator will be oxidized by air rather than by its
electrooxidation at the electrode. It is preferred that the
reaction of the redox mediator with O.sub.2 proceeds more slowly
than the electrooxidation of the mediator, because if the air
oxidation of the mediator were fast, then dissolved air and the
in-diffusion of air might affect the outcome of the
measurement.
Because typically the assay takes about 10 minutes or less,
preferably 5 minutes or less, and most preferably about 1 minute or
less, it is preferred that the mediator, though air oxidizable in
storage, will not be oxidized by dissolved oxygen during the time
of the assay. Thus, mediators that are not air oxidized in 1
minute, and preferably not even in 10 minutes when dissolved in
plasma or in serum, are preferred. Typically, less than 5%, and
preferably less than 1%, of the reduced mediator should be oxidized
by air during an assay.
The reaction rate of the air oxidation of the mediator cart be
controlled through choice of an appropriate complexing polymer. For
example, the oxidation reaction is much faster for
[Os(4,4'-dimethoxy-2,2'-bipyridine).sub.2Cl].sup.+/+2
coordinatively coupled to poly(1-vinyl imidazole) than for the same
Os complex coupled to poly(4-vinyl pyridine). The choice of an
appropriate polymer will depend on the expected analyte
concentration and the potential applied between the electrodes,
both of which determine the rate of the electrochemical
reaction.
Thus, in one embodiment of the invention, the preferred redox
mediator has the following characteristics: 1) the mediator does
not react with any molecules in the sample or in the sensor other
than the analyte (optionally, via a second electron transfer
agent); 2) nearly all of the redox mediator is oxidized by an
oxidizer such as O.sub.2 prior to introduction of the sample in the
sensor; and 3) the oxidation of the redox mediator by the oxidizer
is slow compared to the electrooxidation of the mediator by the
electrode.
Alternatively, if the redox mediator is to be oxidized in the
presence of the analyte and electroreduced at the electrode, a
reducer rather than an oxidizer would be required. The same
considerations for the appropriate choice of reducer and mediator
apply as described hereinabove for the oxidizer.
The use of stable air-oxidizable redox mediators in the
electrochemical sensors of the invention provides an additional
advantage during storage and packaging. Sensors of the invention
which include air-oxidizable redox mediators can be packaged in an
atmosphere containing molecular oxygen and stored for long periods
of time, e.g., greater than one month, while maintaining at least
80% and preferably at least 90% of the redox species in the
oxidized state.
Use of the Air-Oxidizable Mediators in Optical Sensors
The air-oxidizable redox species of the present invention can be
used in other types of sensors. The osmium complexes described
hereinabove are suitable for use in optical sensors, due to the
difference in the absorption spectra, luminescence and/or
fluorescence characteristics of the complexed Os.sup.+2 and
Os.sup.+3 species. Absorption, transmission, reflection,
luminescence and/or fluorescence measurements of the redox species
will correlate with the amount of analyte in the sample (after
reaction between an analyte and the redox species, either directly,
or via a second electron transfer agent such as an enzyme). In this
configuration, the molar amount of redox mediator should be
greater, on a stoichiometric basis, than the molar amount of
analyte reasonably expected to fill the measurement zone of the
sensor.
Standard optical sensors, including light-guiding optical fiber
sensors, and measurement techniques can be adapted for use with the
air-oxidizable mediators. For example, the optical sensors of the
invention may include a light-transmitting or light reflecting
support on which the air-oxidizable redox species, and preferably
an analyte-responsive enzyme, is coated to form a film. The support
film forms one boundary for the measurement zone in which the
sample is placed. The other boundaries of the measurement zone are
determined by the configuration of the cell. Upon filling the
measurement zone with an analyte-containing sample, reduction of
the air-oxidizable mediator by the analyte, preferably via reaction
with the analyte-responsive enzyme, causes a shift in the
mediator's oxidation state that is detected by a change in the
light transmission, absorption, or reflection spectra or in the
luminescence and/or fluorescence of the mediator at one or more
wavelengths of light.
Multiple Electrode Sensors and Calibration
Multiple electrode sensors may be used for a variety of reasons.
For example, multiple electrode sensors may be used to test a
variety of analytes using a single sample. One embodiment of a
multiple electrode sensor, shown in FIG. 5, has one or more sample
chambers which in turn may contain one or more working electrodes
22 with each working electrode 22 defining a different measurement
zone. If the redox mediator is non-leachable, one or more of the
working electrodes have the appropriate chemical reagents, for
example, an appropriate enzyme, to test a first analyte and one or
more of the remaining working electrodes have appropriate chemical
reagents to test a second analyte. The chemical reagents (e.g.,
redox mediator and/or second electron transfer agent) can be
deposited as a sensing layer on the working electrode or, if
diffusible reagents are used, they can be deposited on any surface
of the sample chamber or placed in the sample. For example, a
multiple electrode sensor might include 1) one or more working
electrodes having glucose oxidase in the sensing layer to determine
glucose concentration and 2) one or more working electrodes having
lactate oxidase in the sensing layer to determine lactate
concentration.
Multiple electrode sensors may also be used to improve the
precision of the resulting readings. The measurements from each of
the working electrodes (all or which are detecting the same
analyte) can be averaged together to obtain a more precise reading.
In some cases, measurements may be rejected if the difference
between the value and the average exceeds a threshold limit. This
threshold limit may be, for example, determined based on a
statistical parameter, such as the standard deviation of the
averaged measurements. The average may then be recalculated while
omitting the rejected values. Furthermore, subsequent readings from
an electrode that produced a rejected value may be ignored in later
tests if it is assumed that the particular electrode is faulty.
Alternatively, a particular electrode may be rejected only after
having a predetermined number of readings rejected based on the
readings from the other electrodes.
In addition to using multiple electrode sensors to increase
precision, multiple measurements may be made at each electrode and
averaged together to increase precision. This technique may also be
used with a single electrode sensor to increase precision.
Errors in assays may occur when mass produced sensors are used
because of variations in the volume of the measurement zone of the
sensors. Two of the three dimensions of the measurement zone, the
length and the width, are usually relatively large, between about
1-5 mm. Electrodes of such dimensions can be readily produced with
a variance of 2% or less. The submicroliter measurement zone volume
requires, however, that the third dimension be smaller than the
length or width by one or two order of magnitude. As mentioned
hereinabove, the thickness of the sample chamber is typically
between about 50 and about 200 .mu.m. Manufacturing variances in
the thickness may be on the order of 20 to 50 .mu.m. Therefore, it
may be desirable that a method be provided to accommodate for this
uncertainty in the volume of sample within the measurement
zone.
In one embodiment of the invention, depicted in FIG. 5, multiple
working electrodes 42, 44, 46 are provided on a base material 48.
These electrodes are covered by another base, not shown, which has
counter electrodes, not shown, disposed upon it to provide multiple
facing electrode pairs. The variance in the separation distance
between the working electrode and the counter electrode among the
electrode pairs on a given sensor is significantly reduced, because
the working electrodes and counter electrodes are each provided on
a single base with the same spacer 28 between each electrode pair
(see FIG. 3).
One example of a multiple electrode sensor that can be used to
accurately determine the volume of the measurement zones of the
electrode pairs and that is also useful in reducing noise is
presented herein. In this example, one of the working electrodes 42
is prepared with a non-leachable redox mediator and a non-leachable
second electron transfer agent (e.g., an enzyme). Sorbent material
may be disposed between that working electrode 42 and its
corresponding counter electrode. Another working electrode 44
includes non-leachable redox mediator, but no second electron
transfer agent on the electrode. Again, this second electrode pair
may have sorbent material between the working electrode 44 and the
corresponding counter electrode. An optional third working
electrode 46 has no redox mediator and no second electron transfer
agent bound to the electrode, nor is there sorbent material between
the working electrode 46 and its corresponding counter electrode. A
similar configuration can be constructed using diffusible redox
mediator and/or diffusible second electron transfer agent although
diffusible components are not limited to being disposed on the
working electrode. In some instances, the distance between
electrode pairs is sufficient that redox mediator and/or enzyme do
not substantially diffuse between electrode pairs within the
measurement period and/or in the time period from introduction of
the same sample into the sample chamber to the end of the
measurement.
The sensor error caused by redox mediator in a non-uniform
oxidation state prior to the introduction of the sample can be
measured by concurrently electrolyzing the sample in the
measurement zones that are proximate electrodes 42 and 44. At
electrode 42, the analyte is electrolyzed to provide the sample
signal. At electrode 44, the analyte is not electrolyzed because of
the absence of the second electron transfer agent (assuming that a
second electron transfer agent is necessary). However, a charge
will pass (and a current will flow) due to the electrolysis of the
redox mediator that was in a mixed oxidation state (i.e., some
redox centers in the reduced state and some in the oxidized state)
prior to the introduction of the sample and/or the shuttling of a
diffusible redox mediator between the working electrode and the
counter electrode. The small charge passed between the electrodes
in this second electrode pair can be subtracted from the charge
passed between the first electrode pair to substantially remove the
error due to the oxidation state of the redox mediator and/or to
remove the background current caused by a diffusible redox
mediator. This procedure also reduces the error associated with
other electrolyzed interferents, such as ascorbate, urate, and
acetaminophen, as well as errors associated with capacitive
charging and faradaic currents.
The thickness of the sample chamber can be determined by measuring
the capacitance, preferably in the absence of any fluid, between
electrode 46 (or any of the other electrodes 42, 44 in the absence
of sorbent material) and its corresponding counter electrode. The
capacitance of an electrode pair depends on the surface area of the
electrodes, the interelectrode spacing, and the dielectric constant
of the material between the plates. The dielectric constant of air
is unity which typically means that the capacitance of this
electrode configuration is a few picofarads (or about 100-1000
picofarads if there is fluid between the electrode and counter
electrode given that the dielectric constant for most biological
fluids is approximately 75). Thus, since the surface area of the
electrodes are known, measurement of the capacitance of the
electrode pair allows for the determination of the thickness of the
measurement zone to within about 1-5%.
The amount of void volume in the sorbent material, can be
determined by measuring the capacitance between electrode 44 (which
has no second electron transfer agent) and its associated counter
electrode, both before and after fluid is added. Upon adding fluid,
the capacitance increases markedly since the fluid has a much
larger dielectric constant. Measuring the capacitance both with and
without fluid allows the determination of the spacing between the
electrodes and the void volume in the sorbent, and thus the volume
of the fluid in the reaction zone.
Other electrode configurations can also use these techniques (i.e.,
capacitance measurements and coulometric measurements in the
absence of a critical component) to reduce background noise and
error due to interferents and imprecise knowledge of the volume of
the interrogated sample. Protocols involving one or more electrode
pairs and one or more of the measurements described above can be
developed and are within the scope of the invention. For example,
only one electrode pair is needed for the capacitance measurements,
however, additional electrode pairs may be used for
convenience.
Fill Indicator
When using a sample chamber that is filled with 1 .mu.L or less of
fluid, it is often desirable to be able to determine when the
sample chamber is filled. FIGS. 18A-18C illustrate one sensor
having a fill indicator structure. FIG. 18A illustrates a first
substrate 500 upon which a working electrode 502 is printed. A
spacer 504 (FIG. 1813), such as, for example, a layer of adhesive
or a double-sided tape, is formed over the first substrate 500 and
working electrode 502 with a channel 506 formed in the layer to
provide a sample chamber. A second substrate 508 is printed with
two counter electrodes 510, 512, as shown in FIG. 18C (inverted
with respect to FIGS. 18A and 18B to show the electrode side up).
In some instances, the counter electrode 510 nearest an entrance
514 of the channel 506 has a surface area within the sample chamber
that is at least two times larger than the other counter electrode
512, and preferably at least five or ten times larger.
The sensor can be indicated as filled by observing a signal between
the second counter electrode 512 and the working electrode 502 as
the sensor fills with fluid. When fluid reaches the second counter
electrode 512, the signal from that counter electrode should
change. Suitable signals for observing include, for example,
voltage, current, resistance, impedance, or capacitance between the
second counter electrode 512 and the working electrode 502.
Alternatively, the sensor may be observed after filling to
determine if a value of the signal (e.g., voltage, current,
resistance, impedance, or capacitance) has been reached indicating
that the sample chamber is filled.
In alternative embodiments, the counter electrode and/or working
electrode may be divided into two or more parts and the signals
from the respective parts observed to determine whether the sensor
has been filled. In one example, the working electrode is in a
facing relationship with the counter electrode and the indicator
electrode. In another example, the counter electrode, working
electrode, and indicator electrode are not in a facing
relationship, but may be, for example, side-by-side. In other
cases, a second electrode pair may be used with signals from the
second electrode pair being monitored for changes and/or for
approaching a particular value to determine that the sensor has
filled. Typically, the indicator electrode is further downstream
from a sample inlet port than the working electrode and counter
electrode.
For side-fill sensors, such as those illustrated in FIGS. 19A-19C
and 20A-20C, two indicator electrodes may be disposed on either
side of the primary counter electrode. This permits the user to
fill the sample chamber from either the left or right side with an
indicator electrode disposed further upstream. This three-electrode
configuration is not necessary. Side-fill sensors can also have a
single indicator electrode and, preferably, some indication as to
which side should be placed in contact with the sample fluid.
In one embodiment, the use of three counter/reference electrodes
and/or indicator electrodes, detects when the sample chamber begins
to fill and when the sample chamber has been filled to prevent
partial filling of the sample chamber. In this embodiment, the two
indicator electrodes are held at a different potential than the
largest counter/reference electrode. The start and completion of
filling of the sample chamber is indicated by the flow of current
between the indicator and counter/reference electrodes.
In other instances, the potential of each of the counter/reference
electrodes may be the same. When the potential at all three
counter/reference electrodes is the same for example, 0 volts, then
as the measurement zone begins to fill, the fluid allows for
electrical contact between a working electrode and the first
counter/reference electrode, causing a current at the first
counter/reference electrode due to the reaction of the analyte with
the enzyme and the mediator. When the fluid reaches the third
counter/reference electrode, another current may be measured
similar to the first counter/reference electrode indicating that
the measurement zone is full. When the measurement zone is full,
the three counter/reference electrodes may be shorted together or
their signals may be added or otherwise combined.
The indicator electrode may also be used to improve the precision
of the analyte measurements according to the methods described
above for multiple electrode sensors. The indicator electrode may
operate as a working electrode or as a counter electrode or
counter/reference electrode. In the embodiment of FIGS. 18A-18C,
the indicator electrode 512 can act as a second counter or
counter/reference electrode with respect to the working electrode
502. Measurements from the indicator electrode/working electrode
pair can be combined (for example, added to and/or averaged) with
those from the first counter or counter/reference electrode/working
electrode pair to obtain more accurate measurements. In one
embodiment, the indicator electrode may operate as a second working
electrode with the counter electrode or counter/reference
electrode. In another embodiment, the indicator electrode may
operate as a second working electrode with a second counter
electrode or counter/reference electrode. In still another
embodiment, the indicator electrode may operate as a second counter
electrode or counter/reference electrode with a second working
electrode.
The sensor or a sensor reader may include a sign (e.g., a visual
sign or auditory signal) that is activated in response to the
indicator electrode to alert the user that the measurement zone has
been filled. In some instances, the sensor or a sensor reader may
be configured to initiate a reading when the indicator electrode
indicates that the measurement zone has been filled with or without
alerting the user. The reading can be initiated, for example, by
applying a potential between the working electrode and the counter
electrode and beginning to monitor the signals generated at the
working electrode.
Heating of Sample
The sample may be heated to increase the rate of diffusion,
oxidation, or reduction of the analyte. This heating may be
accomplished by a variety of techniques including placing the
sensor in a heated environment or applying a heating unit to the
sensor.
Another technique includes providing a thermal heating element,
such as, for example, a wire or an ink that is capable of
converting electrical energy into heat energy, on the sensor. This
wire or ink can be applied, for example, on the opposite side of a
base material, such as a polymer film, from one or more of the
working, counter, reference, or counter/reference electrodes, or
applied around the periphery of the working, counter, reference, or
counter/reference electrodes. In some instances, the sample may be
heated up to 5 to 20.degree. C. above an initial temperature. In
other instances, the temperature of the sample may not be known but
a constant amount of power or current may be applied to the wire or
ink.
EXAMPLES
The invention will be further characterized by the following
examples. These examples are not meant to limit the scope of the
invention which has been fully set forth in the foregoing
description. Variations within the concepts of the invention are
apparent to those skilled in the art.
Example 1
Preparation of a Small Volume In Vitro Sensor for the Determination
of Glucose Concentration
A sensor was constructed corresponding to the embodiment of the
invention depicted in FIG. 1. The working electrode was constructed
on a Mylar.TM. film (DuPont), the Mylar.TM. film having a thickness
of 0.175 mm and a diameter of about 2.5 cm. An approximately 12
micron thick carbon pad having a diameter of about 1 cm was screen
printed on the Mylar.TM. film. The carbon electrode was overlaid
with a water-insoluble dielectric insulator (Insulayer) having a
thickness of 12 .mu.m, and a 4 mm diameter opening in the
center.
The center of the carbon electrode, which was not covered by the
dielectric, was coated with a non-leachable redox mediator. The
redox mediator was formed by complexing poly(1-vinyl imidazole)
with Os(4,4'-dimethoxy-2,2'-bipyridine).sub.2Cl.sub.2 followed by
cross-linking glucose oxidase with the osmium polymer using
polyethylene glycol diglycidyl ether (PEGDGE) as described in
Taylor et al., J. Electroanal. Chem., 396:511 (1995). The ratio of
osmium to imidazole functionalities in the redox mediator was
approximately 1:15. The mediator was deposited on the working
electrode in a layer having a thickness of 0.6 .mu.m and a diameter
of 4 mm. The coverage of the mediator on the electrode was about 60
.mu.g/cm.sup.2 (dry weight). A spacer material was placed on the
electrode surrounding the mediator-covered surface of the
electrode. The spacer was made of polytetrafluoroethylene (PTFE)
and had a thickness of about 0.040 mm.
A sorbent material was placed in contact with the mediator-covered
surface of the working electrode. The sorbent was made of nylon
(Tetko Nitex nylon 3-10/2). The sorbent had a diameter of 5 mm, a
thickness of 0.045 mm, and a void volume of about 20%. The volume
of sample in the measurement zone was calculated from the
dimensions and characteristics of the sorbent and the electrode.
The measurement zone had a diameter of 4 mm (the diameter of the
mediator covered surface of the electrode) and a thickness of 0.045
mm (thickness of the nylon sorbent) to give a volume of 0.57 .mu.L.
Of this space, about 80% was filled with nylon and the other 20%
was void space within the nylon sorbent. This resulting volume of
sample within the measurement zone was about 0.11 .mu.L.
A counter/reference electrode was placed in contact with the spacer
and the side of the sorbent opposite to the working electrode so
that the two electrodes were facing each other. The
counter/reference electrode was constructed on a Mylar.TM. film
having a thickness of 0.175 mm and a diameter of about 2.5 cm onto
which a 12 micron thick layer of silver/silver chloride having a
diameter of about 1 cm was screen printed.
The electrodes, sorbent, and spacer were pressed together using
plates on either side of the electrode assembly. The plates were
formed of polycarbonate plastic and were securely clamped to keep
the sensor together. The electrodes were stored in air for 48 hours
prior to use.
Tabs extended from both the working electrode and the
counter/reference electrode and provided for an electrical contact
with the analyzing equipment. A potentiostat was used to apply a
potential difference of +200 mV between the working and
counter/reference electrodes, with the working electrode being the
anode. There was no current flow between the electrodes in the
absence of sample, which was expected, as no conductive path
between the electrodes was present.
The sample was introduced via a small tab of nylon sorbent material
formed as an extension from the nylon sorbent in the sample
chamber. Liquid was wicked into the sorbent when contact was made
between the sample and the sorbent tab. As the sample chamber
filled and the sample made contact with the electrodes, current
flowed between the electrodes. When glucose molecules in the sample
came in contact with the glucose oxidase on the working electrode,
the glucose molecules were electrooxidized to gluconolactone. The
osmium redox centers in the redox mediator then reoxidized the
glucose oxidase. The osmium centers were in turn reoxidized by
reaction with the working electrode. This provided a current which
was measured and simultaneously integrated by a coulometer
(EG&G Princeton Applied Research Model #173).
The electrochemical reaction continued until the current reached a
steady state value which indicated that greater than 95% of the
glucose had been electroreduced. The current curve obtained by
measurement of the current at specific intervals was integrated to
determine the amount of charge passed during the electrochemical
reaction. These charges were then plotted versus the known glucose
concentration to produce a calibration curve.
The sensor was tested using 0.5 .mu.L aliquots of solutions
containing known concentrations of glucose in a buffer of
artificial cerebrospinal fluid or in a control serum (Baxter-Dade,
Monitrol Level 1, Miami, Fla.) in the range of 3 to 20 mM glucose.
The artificial cerebrospinal fluid was prepared as a mixture of the
following salts: 126 mM NaCl, 27.5 mM NaHCO.sub.3, 2.4 mM KCl, 0.5
mM KH.sub.2PO.sub.4, 1.1 mM CaCl.sub.2.2H.sub.2O, and 0.5 mM
Na.sub.2SO.sub.4.
The results of the analyses are shown in Table 1 and in FIG. 7. In
Table 1, Q.sub.avg is the average charge used to electrolyze the
glucose in 3-6 identical test samples (FIG. 7 graphs the charge for
each of the test samples) and the 90% rise time corresponds to the
amount of time required for 90% of the glucose to be electrolyzed.
The data show a sensor precision of 10-20%, indicating adequate
sensitivity of the sensor for low glucose concentrations, as well
as in the physiologically relevant range (30 .mu.g/dL-600
.mu.g/dL).
TABLE-US-00001 TABLE 1 Sensor Results Using Glucose Oxidase Number
of Samples 90% rise time Tested Q.sub.avg (.mu.C) (sec) buffer only
4 9.9 .+-. 1.8 13 .+-. 6 3 mM glucose/buffer 5 17.8 .+-. 3.5 19
.+-. 5 6 mM glucose/buffer 4 49.4 .+-. 4.9 25 .+-. 3 10 mM
glucose/buffer 6 96.1 .+-. 12.4 36 .+-. 17 15 mM glucose/buffer 5
205.2 .+-. 75.7 56 .+-. 23 20 mM glucose/buffer 4 255.7 .+-. 41.0
62 .+-. 17 4.2 mM glucose/serum 3 44.2 .+-. 4.3 44 .+-. 3 15.8 mM
glucose/serum 3 218.2 .+-. 57.5 72 .+-. 21
The average measured values of glucose concentration were fit by
one or more equations to provide a calibration curve. FIG. 8 shows
the calibration curves for the glucose/buffer data of Table 1. One
of the 15.0 mM glucose measurements was omitted from these
calculations because it was more than two standard deviations away
from the average of the measurements. The higher glucose
concentrations (10-20 mM) were fit by a linear equation. The lower
glucose concentrations were fit by a second order polynomial.
FIG. 9 shows the data of Table 1 plotted on an error grid developed
by Clarke et al., Diabetes Care, 5, 622-27, 1987, for the
determination of the outcome of errors based on inaccurate glucose
concentration determination. The graph plots "true" glucose
concentration vs. measured glucose concentration, where the
measured glucose concentration is determined by calculating a
glucose concentration using the calibration curves of FIG. 8 for
each data point of FIG. 7. Points in zone A are accurate, those in
zone B are clinically acceptable, and those in zones C, D, and E
lead to increasingly inappropriate and finally dangerous
treatments.
There were 34 data points. Of those data points 91% fell in zone A,
6% in zone B, and 3% in zone C. Only one reading was determined to
be in zone C. This reading was off-scale and is not shown in FIG.
9. Thus, 97% of the readings fell in the clinically acceptable
zones A and B.
The total number of Os atoms was determined by reducing all of the
Os and then electrooxidizing it with a glucose-free buffer in the
sample chamber. This resulted in a charge of 59.6+5.4 .mu.C.
Comparison of this result with the glucose-free buffer result in
Table 1 indicated that less than 20% of the Os is in the reduced
form prior to introduction of the sample. The variability in the
quantity of osmium in the reduced state is less than 5% of the
total quantity of osmium present.
Example 2
Response of the Glucose Sensor to Interferents
A sensor constructed in the same manner as described above for
Example 1 was used to determine the sensor's response to
interferents. The primary electrochemical interferents for blood
glucose measurements are ascorbate, acetaminophen, and urate. The
normal physiological or therapeutic (in the case of acetaminophen)
concentration ranges of these common interferents are:
ascorbate: 0.034-0.114 mM
acetaminophen: 0.066-0.200 mM
urate (adult male): 0.27-0.47 mM
Tietz, in: Textbook of Clinical Chemistry, C. A. Burtis and E. R.
Ashwood, eds., W.B. Saunders Co., Philadelphia 1994, pp.
2210-12.
Buffered glucose-free interferent solutions were tested with
concentrations of the interferents at the high end of the
physiological or therapeutic ranges listed above. The injected
sample volume in each case was 0.5 .mu.L. A potential of +100 mV or
+200 mV was applied between the electrodes. The average charge
(Q.sub.avg) was calculated by subtracting an average background
current obtained from a buffer-only (i.e., interferent-free)
solution from an average signal recorded with interferents present.
The resulting average charge was compared with the signals from
Table 1 for 4 mM and 10 mM glucose concentrations to determine the
percent error that would result from the interferent.
TABLE-US-00002 TABLE 2 Interferent Response of Glucose Sensors
Error @ Error @ 4 mM 10 mM Solution E (mV) n Q.sub.avg (.mu.C)
glucose glucose 0.114 mM ascorbate 100 4 0.4 2% <1% 0.114 mM
ascorbate 200 4 -0.5 2% <1% 0.2 mM acetaminophen 100 4 0.1
<1% <1% 0.2 mM acetaminophen 200 4 1.0 5% 1% 0.47 mM urate
100 4 6.0 30% 7% 0.47 mM urate 200 4 18.0 90% 21%
These results indicated that ascorbate and acetaminophen were not
significant interferents for the glucose sensor configuration,
especially for low potential measurements. However, urate provided
significant interference. This interference can be minimized by
calibrating the sensor response to a urate concentration of 0.37
mM, e.g., by subtracting an appropriate amount of charge as
determined by extrapolation from these results from all glucose
measurements of the sensor. The resulting error due to a 0.10 mM
variation in urate concentration (the range of urate concentration
is 0.27-0.47 in an adult male) would be about 6% at 4 mM glucose
and 100 mV.
Example 3
Sensor with Glucose Dehydrogenase
A sensor similar to that described for Example 1 was prepared and
used for this example, except that glucose oxidase was replaced by
pyrroloquinoline quinone glucose dehydrogenase and a potential of
only +100 mV was applied as opposed to the +200 mV potential in
Example 1. The results are presented in Table 3 below and graphed
in FIG. 10.
TABLE-US-00003 TABLE 3 Sensor Results Using Glucose Dehydrogenase
90% rise time n Q.sub.avg (.mu.C) (s) buffer 4 21.7 .+-. 5.2 14
.+-. 3 3 mM glucose/buffer 4 96.9 .+-. 15.0 24 .+-. 6 6 mM
glucose/buffer 4 190.6 .+-. 18.4 26 .+-. 6 10 mM glucose/buffer 4
327.8 .+-. 69.3 42 .+-. 9
The results indicated that the charge obtained from the glucose
dehydrogenase sensor was much larger than for the comparable
glucose oxidase sensor, especially for low concentrations of
glucose. For 4 mM glucose concentrations the measurements obtained
by the two sensors differed by a factor of five. In addition, the
glucose dehydrogenase sensor operated at a lower potential, thereby
reducing the effects of interferent reactions.
In addition, the results from Table 3 were all fit by a linear
calibration curve as opposed to the results in Example 1, as shown
in FIG. 10. A single linear calibration curve is greatly preferred
to simplify sensor construction and operation.
Also, assuming that the interferent results from Table 2 are
applicable for this sensor, all of the interferents would introduce
an error of less than 7% for a 3 mM glucose solution at a potential
of 100 mV.
Example 4
Determination of Lactate Concentration in a Fluid Stream
The sensor of this Example was constructed using a flow cell
(BioAnalytical Systems, Inc. #MF-1025) with a glassy carbon
electrode. A redox mediator was coated on the electrode of the flow
cell to provide a working electrode. In this case, the redox
mediator was a polymer formed by complexing poly(1-vinyl imidazole)
with Os(4,4'-dimethyl-2,2'-bipyridine).sub.2Cl.sub.2 with a ratio
of 1 osmium for every 15 imidazole functionalitics. Lactate oxidase
was cross-linked with the polymer via polyethylene glycol
diglycidyl ether. The mediator was coated onto the electrode with a
coverage of 500 .mu.g/cm.sup.2 and a thickness of 5 .mu.m. The
mediator was covered by a polycarbonate track-etched membrane
(Osmonics-Poretics #10550) to improve adherence in the flow stream.
The membrane was then overlaid by a single 50 .mu.m thick spacer
gasket (BioAnalytical Systems, Inc. #MF-1062) containing a void
which defined the sample chamber and corresponding measurement
zone. Assembly of the sensor was completed by attachment of a cell
block (BioAnalytical Systems, Inc. #MF-1005) containing the
reference and auxiliary electrodes of the flow cell.
The sample chamber in this case corresponded to a 50 .mu.m thick
cylinder (the thickness of the spacer gasket) in contact with a
mediator-coated electrode having a surface area of 0.031 cm.sup.2.
The calculated volume of sample in the measurement zone of this
sensor was approximately 0.16 .mu.L.
The flow rate of the fluid stream was 5 .mu.L/min. A standard three
electrode potentiostat was attached to the cell leads and a
potential of +200 mV was applied between the redox mediator-coated
glassy carbon electrode and the reference electrode. This potential
was sufficient to drive the enzyme-mediated oxidation of
lactate.
As the fluid stream flowed through the sensor, a steady-state
current proportional to the lactate concentration was measured. At
periodic intervals the fluid flow was stopped and current was
allowed to flow between the electrodes until approximately all of
the lactate in the measurement zone was electrooxidized, as
indicated by the achievement of a stabilized, steady-state current.
The total charge, Q, required for lactate electrooxidation was
found by integration of the differential current registered from
the flow stoppage until the current reached a steady-state. The
concentration was then calculated by the following equation:
[lactate]=Q/2FV (4) where V is the volume of sample within the
measurement zone and F is Faraday's constant.
This assay was performed using lactate solutions having nominal
lactate concentrations of 1.0, 5.0, and 10.0 mM. The measured
concentrations for the assay were 1.9, 5.4, and 8.9 mM
respectively.
Example 5
Determination of the Oxidation State of
Os(4,4'-dimethoxy-2,2'-bipyridine).sub.2Cl.sup.+/+2 Complexed with
poly(1-vinyl imidazole)
A sensor having a three electrode design was commercially obtained
from Ecossensors Ltd., Long Hanborough, England, under the model
name "large area disposable electrode". The sensor contained
parallel and coplanar working, reference and counter electrodes.
The working surface area (0.2 cm.sup.2) and counter electrodes were
formed of printed carbon and the reference electrode was formed of
printed Ag/AgCl. A redox mediator was coated on the carbon working
electrode. The redox mediator was formed by complexation of
poly(1-vinyl imidazole) with
Os(4,4'-dimethoxy-2,2'-bipyridine).sub.2Cl.sub.2 in a ratio of 15
imidazole groups per Os cation followed by cross linking the osmium
polymer with glucose oxidase using polyethylene glycol diglycidyl
ether.
The electrode was cured at room temperature for 24 hours. The
coplanar electrode array was then immersed in a buffered
electrolyte solution, and a potential of 1200 mV (sufficient for
conversion of Os(II) to Os(III),) was applied between the working
electrode and the reference electrode.
Upon application of the potential, an undetectable charge of less
than 1 .mu.C was passed. Subsequent reduction and reoxidation of
the redox mediator yielded a charge for conversion of all Os from
Os(II) to Os(III) of 65 .mu.C. Therefore, more than 98% of the Os
cations in the redox mediator were in the desired oxidized Os(III)
state.
Example 6
Determination of the Oxidation State of the
Os(4,4'-dimethoxy-2,2'-bipyridine).sub.2Cl.sup.+/+2 Complexed with
poly(4-vinyl pyridine)
A similar experiment to that of Example 5 was conducted with the
same working/counter/reference electrode configuration except that
the redox mediator on the working electrode was changed to a
complex of Os(4,4'-dimethoxy-2,2'-bipyridine).sub.2Cl.sub.2 with
poly(4-vinyl pyridine), with 12 pyridine groups per Os cation,
cross linked with glucose oxidase via polyethylene glycol
diglycidyl ether.
Two sensors were constructed. The electrodes of the two sensors
were cured at room temperature for 24 hours. The electrodes were
then immersed in a buffered electrolyte solution and a potential of
+200 mV was applied between the working and reference
electrodes.
Upon application of the potential to the electrodes, a charge of
2.5 .mu.C and 3.8 .mu.C was passed in the two sensors,
respectively. Subsequent reduction and reoxidation of the redox
mediators yielded oxidation charges of 27.9 .mu.C and 28.0 .mu.C,
respectively. Therefore, the sensors originally contained 91% and
86% of the Os cations in the desirable oxidized Os(III) state.
Example 7
Optical Sensor
An optical sensor is constructed by applying a film of redox
polymer with crosslinked enzyme onto a light-transparent support
such as a glass slide. The quantity of redox mediator is equal to
or greater than (in a stoichiometric sense) the maximum quantity of
analyte expected to fill the measurement zone. The spacer material,
sorbent and facing support are securely clamped. The sample chamber
is adapted to transmit light through the assembled sensor to an
optical density detector or to a luminescence and/or fluorescence
detector. As sample fills the sample chamber and the redox mediator
is oxidized, changes in the absorption, transmission, reflection or
luminescence and/or fluorescence of the redox mediator in the
chamber are correlated to the amount of glucose in the sample.
Example 8
Blood Volumes from Upper Arm Lancet Sticks
The forearm of a single individual was pierced with a lancet
multiple times in order to determine the reproducibility of blood
volumes obtained by this method. Despite more than thirty lancet
sticks in the anterior portion of each forearm and the dorsal
region of the left forearm, the individual identified each stick as
virtually painless.
The forearm was pierced with a Payless Color Lancet. The blood from
each stick was collected using a 1 .mu.L capillary tube, and the
volume was determined by measuring the length of the blood column.
The volumes obtained from each stick are shown below in Table
4.
TABLE-US-00004 TABLE 4 Volume of Lancet Sticks Left Anterior Right
Anterior Left Dorsal Forearm, (nL) Forearm, (nL) Forearm, (nL) 1
180 190 180 2 250 180 300 3 170 120 310 4 150 100 300 5 100 210 60
6 50 140 380 7 90 120 220 8 130 140 200 9 120 100 380 10 100 320 11
260 12 250 13 280 14 260 Avg. 138 .+-. 58 nL 140 .+-. 40 nL 264
.+-. 83 nL
Example 9
A Sensor with Diffusible Redox Mediator
A sensor was formed by printing graphite ink (Graphite #G4491,
Ercon, Wareham, Mass.) on a polyester substrate. A mixture of 5.5
.mu.g/cm.sup.2
[Os(dimethyoxybipyridine).sub.2(vinylimidazole)Cl]Cl, 23.7
.mu.g/cm.sup.2 PQQ-glucose dehydrogenase, and 18.2 .mu.g/cm.sup.2
Zonyl FSO.RTM. surfactant (E. I. duPont de Nemours & Co., Inc.,
Wilmington, Del.) were deposited on a portion of the working
electrode. A 150 .mu.m thick pressure sensitive adhesive tape was
then applied to the working electrode leaving only a portion of the
working electrode exposed to form a sample chamber. A second
polyester film with a counter electrode disposed on the film was
provided over the pressure sensitive adhesive tape. The counter
electrode was formed by disposing Ag/AgCl ink (Silver/Silver
Chloride #R414, Ercon, Wareham, Mass.) over the second polyester
film. The Ag/AgCl counter electrode was coated with approximately
100 .mu.g/cm.sup.2 of methylated poly(vinylimidazole) crosslinked
using PEGDGE.
Example 10
Measuring Glucose using Sensor with Diffusible Redox Mediator at a
Potential of 0 V
Sensors were formed as described in Example 9 and used to measure
glucose/buffer solutions at 0, 90, 180, 270, and 360 mg/dL glucose
concentration. The charge measured over time for each of these
solutions is graphed in FIG. 15. In the absence of glucose, the
sensor indicates about 3 mg/dL glucose concentration. FIG. 16
illustrates the measured charge versus glucose concentration for
three sensors at each glucose concentration. The measured charge
varies linearly with glucose concentration similar to what is
observed for sensors using non-leachable redox mediator.
Example 11
Other Sensors Formed Using Diffusible Redox Mediator
Sensors A and B were formed by printing graphite ink (Graphite
#G4491, Ercon, Wareham, Mass.) on a polyester substrate. For Sensor
A, a mixture of 8.0 .mu.g/cm.sup.2
[Os(dimethyoxybipyridine).sub.2(vinyl imidazole)Cl]Cl, 34.7
.mu.g/cm.sup.2 PQQ-glucose dehydrogenase, and 26.6 .mu.g/cm.sup.2
Zonyl FSO.RTM. surfactant (E.I. duPont de Nemours & Co., Inc.,
Wilmington, Del.) were deposited on a portion of the working
electrode. For Sensor B, a mixture of 24 .mu.g/cm.sup.2
[Os(dimethyoxybipyridine).sub.2 (vinyl imidazole)Cl]Cl, 104
.mu.g/cm.sup.2 PQQ-glucose dehydrogenase, and 80 .mu.g/cm.sup.2
Zonyl FSO.RTM. surfactant (E.I. duPont de Nemours & Co., Inc.,
Wilmington, Del.) were deposited on a portion of the working
electrode. A 200 .mu.m pressure sensitive adhesive tape was then
formed over the working electrode of each sensor leaving only a
portion of the working electrode exposed to form a sample chamber.
A second polyester film with a counter electrode disposed on the
film was provided over the pressure sensitive adhesive tape. The
counter electrode of each sensor was formed by disposing Ag/AgCl
ink (Silver/Silver Chloride #R414, Ercon, Wareham, Mass.) over the
second polyester film. The Ag/AgCl counter electrode was coated
with approximately 100 .mu.g/cm.sup.2 of methylated
poly(vinylimidazole) crosslinked using PEGDGE.
Example 12
Varying the Amount of Diffusible Redox Mediator in the Sensor
Sensors A and B were each tested to determine the amount of time
required for electrolysis of the analyte. FIG. 17 illustrates the
results. Increasing the amount of diffusible redox mediator in the
sample decreases the response time of the sensor.
Example 13
Clinical Accuracy of the Small Volume Sensor The sensor of this
Example was constructed corresponding to the embodiment of the
invention depicted in FIGS. 24A, 24B, and 24C. The carbon working
electrode was printed on a Melinex.TM. polyester film (DuPont,
Wilmington, Del.), as described in Example 11. The carbon electrode
was coated with 18 .mu.g/cm.sup.2 Os[(MeO).sub.2bpy].sub.2(1-vinyl
imidazole)Cl.sub.3, 162 .mu.g/cm.sup.2 GDH (Toyobo, Japan), 1.35
.mu.g/cm.sup.2 PQQ (Fluka, Mila, Wis.), and 60 .mu.g/cm.sup.2 Zonyl
FSO (DuPont, Wilmington, Del.). The coatings were applied to the
working electrode at 18.degree. C. and in 50% relative humidity. An
adhesive (50 .mu.m thickness) was placed on the carbon electrode
surrounding the coated surface and forming a channel having a width
of about 0.04 inches.
Two Ag/AgCl counter/reference electrodes were printed on a second
Melinex.TM. polymer film, as described in Example 11. The film was
then brought into contact with the adhesive and the working
electrode film so that the working electrode and two counter
electrodes were facing each other. The counter/reference electrodes
were coated with 142 .mu.g/cm.sup.2 methylated polyvinyl imidazole,
18 .mu.g/cm.sup.2 PEGDGE (PolySciences, Warington, Pa.), and 7
.mu.g/cm.sup.2 Zonyl FSO (DuPont, Wilmington, Del.). One of the
counter electrodes, upstream of the other counter electrode, was
used as an indicator electrode to determine when the sample chamber
was full. The sensors were laminated by three passes with a hand
roller and aged for three days at room temperature over
CaSO.sub.4.
The sensors were constructed so that when sufficient current flowed
between indicator and counter/reference electrodes, an external
circuit emitted a visual signal indicating that the channel
overlying the working electrode was full of blood.
A few days prior to using the sensors, dry capacitance measurements
were taken to determine the uniformity of the sample chamber
volume. The variation in capacitance reflected misalignment of
electrodes and/or variation in adhesive thickness. The mean
capacitance measured was 7.49 pF with a standard deviation of 0.28
pF or 3.8%. The maximum capacitance measured was 8.15 pF and the
minimum capacitance measured was 6.66 pF.
The sensors were used to determine the glucose concentration in
blood samples obtained from 23 people. In the study, the people
ranged from 26 to 76 years of age, fourteen were men, and nine were
women. Six of the people were diagnosed with Type 1 diabetes,
sixteen were diagnosed with Type 2 diabetes, and one person was
unknown regarding diabetic status. The people studied had an
average hematocrit of 40.7% with a standard deviation of 3.9%. The
maximum hematocrit was 49% and the minimum hematocrit was
33.2%.
One blood sample for each person was collected by pricking the
finger of the subject. A small volume sensor was filled with this
residual blood.
Three blood samples for each person were then collected in small
volume sensors by using a 2 mm Carelet.TM. to lance the arm. If an
adequate sample was not obtained in 10 seconds, the area around the
puncture wound was kneaded, and then the sensor was filled. Sixteen
of the sixty-nine samples required that the wound be kneaded.
Three blood samples per person were collected by venipuncture. YSI
blood glucose measurements and hematocrit measurements were taken
on at least one sample. Forty-six small volume sensors were also
filled with blood from these samples.
Measurements from the sensor were performed at an applied potential
of 0 mV. BAS potentiostats (CV50W, West Lafayette, Ind.) were "on"
before any sample was applied, so that as the strips filled,
electrolysis was immediate. Current collection was for 150 seconds
(this charge is termed "complete" electrolysis), although most
assays were essentially complete well before 150 seconds. No
results were discarded. Three successive sensor blood glucose
measurements were taken.
Measurements for the control samples were performed using YSI blood
glucose measurement (Yellow Springs Instruments, Model 2300
Chemical Glucose Analyzer).
The data was plotted against YSI venous results and a linear
function was determined from the data. All data was collected from
"complete" (150 second) electrolysis of glucose in the sensor.
FIG. 28 shows the data for 69 small volume sensors tested on blood
obtained from the arm. R.sup.2 was 0.947, the average CV
(coefficient of variation) was 4.8%, and the RMS (root mean square)
CV was 5.7%.
FIG. 29 shows the data for 23 small volume sensors tested on blood
obtained from the finger. R.sup.2 was 0.986.
FIG. 30 shows the data for 46 small volume sensors tested on venous
blood. R.sup.2 was 0.986. The average CV was 3.8%. The RMS CV was
4.6%.
The invention has been described with reference to various specific
and preferred embodiments and techniques. However, it will be
apparent to one of ordinarily skill in the art that many variations
and modifications may be made while remaining within the spirit and
scope of the invention.
All publications and patent applications in this specification are
indicative of the level of ordinary skill in the art to which this
invention pertains. All publications and patent applications are
herein incorporated by reference to the same extent as if each
individual publication or patent application was specifically and
individually incorporated by reference.
* * * * *
References